Large-pitch coil configurations for a medical device

ABSTRACT

Techniques related to coils for medical device are disclosed. One example coil may comprise multiple filars, each being formed of a biocompatible beta titanium alloy having an elastic modulus ranging from 30 GigaPascals (GPa) to 90 GPa and comprising at least two elements from a group consisting of titanium, molybdenum, niobium, tantalum, zirconium, chromium, iron and tin. At least one of the multiple filars may be electrically insulated one from another. A structural body, such as a lead body, may carry the coil. One or more filars may carry a low-resistance core.

FIELD

This disclosure relates to a medical apparatus and more particularly toa coil configurations for a medical apparatus such as a lead or leadextension.

BACKGROUND

The medical device industry produces a wide variety of electronic andmechanical devices for treating patient medical conditions such aspacemakers, defibrillators, neurostimulators and therapeutic substancedelivery pumps. Medical devices can be surgically implanted or connectedexternally to the patient receiving treatment. Clinicians use medicaldevices alone or in combination with therapeutic substance therapies andsurgery to treat patient medical conditions. For some medicalconditions, medical devices provide the best and sometimes the onlytherapy to restore an individual to a more healthful condition. One typeof medical device is an implantable stimulation system that can be usedto treat conditions including, but not limited to, pain, movementdisorders, pelvic floor disorders, gastroparesis, and a wide variety ofother medical conditions. Such a system may be connected to astimulation lead with or without extension. The lead may carry one ormore elements such as electrodes and/or other sensors that may beelectrically coupled to the system to deliver electrical stimulationand/or to sense signals from the patient's body. These elements may beelectrically coupled to the system via one or more wires configured ascoils or cables, for instance.

SUMMARY

Techniques related to coils for medical devices, including implantablemedical devices, are disclosed. One example coil may comprise multiplefilars, each being formed of a biocompatible beta titanium alloy havingan elastic modulus ranging from 30 GigaPascals (GPa) to 90 GPa andcomprising at least two elements from a group consisting of titanium,molybdenum, niobium, tantalum, zirconium, chromium, iron and tin. Atleast one of the multiple filars may be electrically insulated one fromanother. A structural body, such as a lead body, may carry the coil.

In one example, the coil may have a large pitch and large filar count.For instance the coil may have twelve or more filars with a pitch of upto 0.10 inches. In a particular embodiment, a coil having twelve filarswith a pitch of 0.08 inches is disclosed. This coil may have an outerdiameter of less than 0.03 inches. In such examples, a ratio of pitch ofthe coil to outer diameter of the coil may be relatively large (e.g.,this ratio may be “three” in some embodiments.) Such a coil has arelatively small profile while being capable of transmitting a largenumber of signals. An example coil with twelve filars may transmit up totwelve signals simultaneously. Moreover, because of the low modulus ofthe beta titanium alloy, the coil will be flexible and exhibit arelatively small bend radius at yield. This allows the coil to be moreeasily navigated to a target therapy location within a body of a patientand also enhances patient comfort.

In one embodiment, at least one of the filars of the coil may include alow-resistance core. The core may be formed of a material having aresistivity of less than 25 micro-ohm-cm. The material forming the coremay be one of silver, tantalum, a tantalum alloy, niobium, a niobiumalloy, platinum, a platinum alloy, palladium, or a palladium alloy.Inclusion of a core within a filar allows the resistance of the filar tobe tuned. For instance, tuning may be accomplished by selecting theratio of a cross-sectional area of the core to a cross-sectional area ofthe filar. By tuning the resistance of the filar to be the same as, orsimilar to, a resistance of an element to which the filar isinterconnected, a quality of signal transmission can be improved. Forinstance, there may be fewer reflections.

Another example of the disclosure provides a method comprising winding acoil of multiple filars, each being formed of a biocompatible betatitanium alloy having an elastic modulus ranging from 30 GigaPascals(GPa) to 90 GPa and comprising at least two elements from a groupconsisting of titanium, molybdenum, niobium, tantalum, zirconium,chromium, iron and tin. At least one of the filars is provided with aninsulating layer to electrically insulate the filar from the otherfilars. Each of the filars may have an insulating layer such that allfilars are insulated from all other filars. A ratio of pitch of the coilto outer diameter of the coil may be greater than one.

The foregoing method may further comprise incorporating the coil into amedical device adapted to perform at least one of delivering therapy toa patient or sensing a signal from the patient. The method may alsoinclude electrically coupling one or more of the filars each to adifferent respective one of multiple elements (e.g., such as electrodes)that are carried by the medical device. The method may also involvetuning a resistance of one or more of the coil filars so that each suchfilar has substantially a same resistance as a resistance of therespective element to which the filar is electrically coupled, asdescribed above.

Another example according to the disclosure includes a medicalelectrical lead. The lead may include a lead body and a coil comprisingmultiple filars, each being formed of a biocompatible beta titaniumalloy having an elastic modulus ranging from 30 GigaPascals (GPa) to 90GPa and comprising at least two elements from a group consisting oftitanium, molybdenum, niobium, tantalum, zirconium, chromium, iron andtin. At least one of the multiple filars may be electrically insulatedfrom the other filars. The coil may contain at twelve filars and have anouter diameter of less than 0.03 inches. The pitch of the coil may besubstantially about 0.08 inches.

Other aspects of the disclosure will become apparent to those skilled inthe art from the following description and the accompanying drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 shows one example system that may usefully employ the techniquesand mechanisms of the current disclosure.

FIG. 2 shows an example implantable neurostimulation system.

FIG. 3A is a side view of one embodiment of a stimulation lead.

FIG. 3B is a cross-sectional view of an array of segmented electrodes.

FIG. 4A is a cross-sectional view of one embodiment of a wire that maybe used to electrically couple one or more conducting electrodes to oneor more connector electrodes.

FIG. 4B is a side cutaway view of a distal end of a lead carrying a betatitanium wire having a low-resistance core electrically and mechanicallycoupled to an electrode.

FIG. 5 is a diagram illustrating S-N curves for Ti-15Mo, MP35N andlow-titanium MP35N (LT MP35N).

FIG. 6 is a graph illustrating resistivities of various alloys.

FIG. 7 is a graph illustrating tunability of resistance for asilver-cored MP35N wire as compared to that for a Ti-15Mo wire having atantalum core.

FIG. 8 is a diagram plotting resistivity against the elastic modulus (E)for various materials.

FIG. 9 is a graph of the level of biocompatibility of various elementsas compared to polarization resistance.

FIG. 10A is a side view of one embodiment of a coil formed of at leastone filar having a beta titanium outer layer and a low-resistance core.

FIG. 10B is a side view of a coil according to another embodiment of thecurrent disclosure.

FIG. 10C is a conceptual drawing of a side cut-away view of a coilcontaining eight filars, each being coupled to a respective one of eightelectrodes.

FIG. 11A is a graph comparing the bend radius to the maximum pitch for afirst coil formed of Ti-15Mo and a second coil formed of MP35N.

FIG. 11B is a graph comparing the bending radius at yield to coil outerdiameter for a four-filar MP35N coil and a four-filar Ti-15Mo coil.

FIG. 11C is a graph comparing the bending radius at yield to coil outerdiameter for an eight-filar MP35N coil and an eight-filar Ti-15Mo coil.

FIG. 12 is a graph comparing the coil pitch to the total length of wirerequired to form a coil.

FIG. 13 is a graph illustrating an enlarged strain plot for Ti-15Mo andMP35N coils having the same outer diameter.

FIG. 14A is a cross-sectional view of one example of a cable accordingto the current disclosure.

FIG. 14B is a side view of a cable according to one embodiment of thedisclosure.

FIG. 15A is a cross-sectional view of a cable formed of wires that donot have low-resistance cores.

FIG. 15B is a side view of a coil formed of independently-insulatedfilars that do not include low-resistance cores.

FIGS. 15A-15D are cross-sectional views of various cable examples formedof wires that do not have low-resistance cores.

FIG. 15E is a side perspective view of a cable such as shown in thecross-sectional view in FIG. 15A.

FIG. 16A is a cross-sectional view of a medical electrical lead that maycarry multiple cables.

FIG. 16B is a cross-sectional view of another example medical electricallead that may carry multiple cables.

FIG. 16C is an example of a medical electrical lead that carries asingle cable.

FIG. 17 is a flow diagram of one method of forming a wire according toone example of the disclosure.

FIG. 18 is a flow diagram of a process according to one specific cableembodiment of the current disclosure.

FIG. 19 is a flow diagram of another process according to one specificcoil embodiment of the current disclosure.

The drawings are not necessarily to scale. Like numbers used in thefigures refer to like components, steps and the like. However, it willbe understood that the use of a number to refer to a component in agiven figure is not intended to limit the component in another figurelabeled with the same number. In addition, the use of different numbersto refer to components is not intended to indicate that the differentnumbered components cannot be the same or similar.

DETAILED DESCRIPTION

FIG. 1 shows one exemplary environmental view 10 of an implantablemedical device (IMD 12 that may usefully employ the techniques andmechanisms of the current disclosure. While this example displays animplantable neurostimulation system embodiment, other medical systems,including cardiac systems and other systems used to delivery therapy to,and/or sense signals from, a living body may employ the conceptsdisclosed herein.

The IMD 12 of the type shown in view 10 may be used to treat conditionssuch as pain, movement disorders, pelvic floor disorders, gastroparesis,incontinence, sexual disfunction, and a wide variety of other medicalconditions. The IMD 12 is typically implanted subcutaneously in thepatient's body 20 at a location selected by the clinician. A stimulationlead 14, which may be coupled to IMD 12 via a lead extension 18, may befixed in place near the location selected by the clinician using adevice such as an anchor. A programmer 22, which may be a clinician orpatient programmer, may be used to program IMD 12 to deliver electricalstimulation via one or more conducting electrodes located along lead 14.

In FIG. 1, lead 14 is shown specifically providing therapy in thepatient's back region. Such therapy may include delivering spinal cordstimulation (SCS), peripheral nerve stimulation (PNS), and peripheralnerve field stimulation (PNFS), for example, as may be used to treatpain.

In other examples, one or more leads 15 (shown dashed) may by coupled toIMD 12 via a lead extension to deliver therapy to a patient's brain. Forinstance, deep brain stimulation (DBS) or cortical stimulation (CS) maybe used to treat a number of neurological conditions, including, e.g.,depression, dementia, obsessive-compulsive disorder, migraines, eatingdisorders, and movement disorders, such as Parkinson's disease,spasticity, epilepsy, and dystonia. One or more leads such as lead 15may be placed at any location within the brain of the patient, includingthe pedunculopontine nucleus (PPN), thalamus, basal ganglia structures(e.g., globus pallidus, substantia nigra, subthalmic nucleus), zonainserta, fiber tracts, lenticular fasciculus (and branches thereof),ansa lenticularis, and/or the Field of Forel (thalamic fasciculus). Inthe case of migraines, leads 15 may be implanted to provide stimulationto the visual cortex of brain in order to reduce or eliminate migraineheadaches afflicting the patient. Additionally, leads 15 may beimplanted to provide stimulation to the cerebral cortex of brain for thetreatment of epilepsy. The target therapy delivery site may depend uponthe patient condition or disorder being treated. Of course, as discussedabove, leads may extend to other locations in the body beyond what isillustrated in FIG. 1 to treat other medical conditions, and theconditions specifically mentioned above are merely examples.

FIG. 2 shows IMD 12, a stimulation lead 14, and a lead extension 18. TheIMD 12 has a housing 24 (a “can”), a power supply (e.g., a battery orcapacitor) carried in the housing 24, and stimulation electronics whichmay be powered by the power supply. These electronics may be coupled toa connector block 26, which is also known as a terminal block.

FIG. 3A is a side view of one embodiment of stimulation lead 14. Thestimulation lead 14 has a lead proximal end 30, a lead distal end 32 anda structural body 34 which may be formed of a material that is anelectrical insulator in one example. The lead distal end 32 has one ormore conducting electrodes 38 to deliver electrical stimulation to aliving body or to receive an electrical signal from the body. The leadproximal end 30 includes one or more connector electrodes 36 (also knownas electrical terminals) to electrically couple the one or moreconducting electrodes 38 to stimulation electronics inside IMD 12,optionally through lead extension 18.

Using stimulation electronics within IMD 12, any one or more of theconducting electrodes 38 may be activated as cathodes and one or moreothers of these electrodes may be activated as anodes to deliverstimulation to the patient. Alternatively or additionally, an electrodeon the case of IMD 12 may serve as an anode or cathode. In someinstances, some or all of electrodes 38 may be employed to sense signalsfrom a patient's body. In some embodiments, such sensing of signals maybe time-multiplexed with delivery of therapy via these electrodes.

In one embodiment, conducting electrodes 38 may be ring electrodes thatencircle the circumference of the body of lead 14. Alternatively,electrodes having a different geometry may be carried by the lead. Forinstance, segmented electrodes may be employed that extend over aportion of the circumference of the lead body.

FIG. 3B is a cross-sectional view of an example of a stimulation lead 14a carrying at least three segmented electrodes 38 a, 38 b, and 38 c. Thesegmented electrodes may be used to deliver electrical stimulation to amore localized area of tissue. For instance, by activating electrode 38a, tissue adjacent to this electrode may be stimulated while tissueresiding near electrode 38 c may not be stimulated.

In one example, the array of electrodes 38 a, 38 b, and 38 c may besubstituted in place of a single ring electrode. For instance, inreference to lead 14 of FIG. 3A, an array of electrodes 38 a-38 c maysubstituted for each of the four electrodes 38 such that lead 14 carriesa complex array of twelve conducting electrodes along its lead body. Insuch an embodiment, it may be desirable to likewise provide twelveconnector electrodes 36, each to couple a respective one of theconducting electrodes to the stimulation electronics of IMD 12. As inthe scenario described above with respect to FIG. 3A, any one or more ofthese twelve segmented conducting electrodes may be activated ascathodes and one or more other electrodes may be activated as anodes.Alternatively or additionally, an electrode on the can of IMD 12 may beactivated as an anode or cathode. A complex electrode array of this typewill allow for generation of a more complex electrical field and mayprovide more targeted stimulation. This may be beneficial in thetreatment of various conditions, including those associated with pain.As a specific example, this may be beneficial when DBS therapy is beingdelivered, since such an array may provide stimulation to very localizedareas in a patient's brain to optimize treatment and minimize sideeffects.

Within a body of a lead such as lead 14 or lead 14 a, there is at leastone wire (not shown) to electrically connect a conducting electrode(e.g., electrodes 38, 38 a-38 c) to a respective connector electrode(e.g., electrodes 36.) Many factors may be taken into account whendetermining what type of wire should be used for this purpose. On theone hand, it may be desirable to utilize a material possessing a lowresistivity to form the wire. This will result in a minimum voltage dropbetween the conducting electrode 38 and the connector electrode 36 towhich it is coupled as well as minimal power loss within the wire.Silver may be used to form a wire that meets these objectives.

Limitations exist with selecting a low-resistivity material alone toserve as the wire. Materials such as silver do not have tensile strengthand easily oxidize. Therefore, bending and twisting a wire made solelyof silver may result in weakening of the wire so that a lead containingthe wire may require replacement more quickly.

One way to address the foregoing limitations is to surround alow-resistance core (e.g., a silver core) with another material that canprovide additional strength. In the past, MP35N (CoNiCrMo) was used forthis purpose.

FIG. 4A is a cross-sectional view of one embodiment of a wire 39 thatmay be used to electrically couple one or more conducting electrodes 38to one or more connector electrodes 36. The wire 39 includes alow-resistance core 40 that is capable of providing a low-resistant pathbetween a conducting electrode 38 and a connector electrode 36. The core40 is directly surrounded in one embodiment by a tube 42 formed of abiocompatible, corrosion-resistant material having a tensile strengththat is substantially higher than the core. As discussed above, thebiocompatible material used for this purpose has typically been MP35Nwith core 40 being made of silver.

Wire 39 may be electrically insulated by a insulating layer 46, whichmay be a polymer. The polymer could be, but is not limited to, ethylenetetrafluoroethylene (ETFE), polytetrafluoroethylene (PTFE), siliconerubber or polyurethane. Other materials that act as electricalinsulators may be used in the alternative.

A wire 39 of the type shown in FIG. 4A that is made of MP35N having asilver core is difficult to manufacture. During the manufacturingprocess, a tube 42 of MP35N is formed that has a diameter of about 0.25cm-1.00 cm. The low-resistance core 40, which in this example may besilver, is threaded into this tube. The tube 42 and the inner core 40are drawn to obtain a wire having a smaller diameter. The drawn wire maythen be heated and re-drawn to form an even smaller wire. The steps ofdrawing, and then heating, the tube 42 and the core 40 are repeatedmultiple times until a wire having a sufficiently small diameter isobtained.

As discussed above, to draw the tube 42 to form a wire having thedesired diameter, the tube is heated to a desired annealing temperatureat which the material becomes flexible. MP35N must generally be annealedat a relatively high temperature above 1000° C. before it can optimallybe drawn. However, the melting point of low-resistivity materials of thetype used to form core 40 is generally below such high temperatures. Forinstance, the melting point of silver is around 960° C.

In view of the foregoing, at the optimal temperature for heating anddrawing an MP35N tube 42, the core material will be melted. As a result,some of the molten core material may exit (i.e., “run out of”) the endsof the tube 42, making the ends of the resulting wire unusable.Moreover, the liquefied core material that is located within the centeraway from the ends of the tube 42 will expand and exert pressure on theinside of the tube. This creates bulges in the tube 42 before and duringthe drawing process. The portions of the wire 39 having suchimperfections will be discarded after the drawing process is completed,resulting in manufacturing waste and lower yields.

To prevent the foregoing from occurring, heating may be limited tosomething much less than what is optimal for MP35N. For instance,heating of a wire containing an MP35N tube 42 surrounding a silver coremay be limited to something under 960° C. to prevent melting of thesilver core. However, at this temperature, the MP35N is relativelybrittle (has a low ductility), making the material difficult to work,and further complicating the manufacturing process. Moreover, theembrittled MP35N wire cannot withstand the repeated stress and strainsof the type that is present in typical implant scenarios (e.g., as inrepeated flexing of a lead body). Thus, such devices may not have aslong of a life, requiring explant or replacement sooner than wouldotherwise be the case.

Another option for addressing the foregoing challenges may involvethreading core 40 into tube 42 after the tube has been drawn to thedesired length and diameter. However, after the tube has been drawn, thetube will have a very small diameter and a very long length. This makesthe threading process difficult, if not impossible in some cases.

The foregoing challenges are addressed by an improved wire 39 that mayuse a low-resistance core 40 such as silver or another low-resistivitymaterial that may be directly surrounded by a tube 42 of a biocompatiblebeta titanium alloy. As known in the art, beta titanium alloys exhibitthe body center cubic (BCC) structure of titanium. This is in contrastto alpha titanium alloys which exhibit the hexagonal close pack (HCP)form of the element. Biocompatible beta titanium alloys may include as amajor alloy one or more of the elements titanium, molybdenum, niobium,tantalum, zirconium, chromium, iron and tin. Specific examples ofbiocompatible beta titanium alloys that may be used according to thisdisclosure include Ti-15Mo, TiOsteum (Ti-35Nb-7Zr-5Ta), TNTZ(Ti-29Nb-12Ta-5Zr), TNCS (Ti-19Nb-5Cr-4Sn), Ti—Nb—Cr—Zr(Ti-20Nb-5Cr-4Zr), TMFZ, TLM (Ti-22Nb-3Zr-3Mo-2Sn) and Ti-30Ta, althoughthese examples are not limiting. Of course, in addition to the majoralloys listed above, these alloys may include trace amounts of otherelements such as silicon (Si), boron (B), or oxygen (O) which may beadded, in one example, to achieve a smaller grain size.

An insulating layer 46 may be provided surrounding the beta titaniumlayer. The insulating layer 46 may be any of the polymers set forthabove. An example use of such a wire is shown in FIG. 4B and will bediscussed further below.

The following Table 1 provides some examples of biocompatible betatitanium alloys, although these examples should not be consideredlimiting.

TABLE 1 Ti Nb Ta Zr Cr Sn Mo TNTZ Balance 29 12 5 TiOsteum Balance 35 57 TNCZ Balance 20 4 5 TNCS Balance 18.6 4.5 4 Ti—45Nb Balance 45 Ti—30TaBalance 30 Ti—15Mo Balance 15 TLM Balance 22.4 2.8 1.8 2.7

A wire formed of a beta titanium alloy may be readily manufacturedwithout the challenges discussed above associated with the need to use alow annealing temperature. For instance, an unbroken tube may be formedof the selected biocompatible beta titanium alloy. In this example, thetube may have an inner diameter of between 0.25 cm-5.0 cm. Alow-resistance core may be inserted into the tube, with the core beingformed of a material that may have a resistivity of less than 25micro-ohm-cm (μΩ-cm) in one example. In some examples, the resistivityof the core material may be between 10 micro-ohm-cm and 20micro-ohms-cm. Examples of materials that may be used to formlow-resistance cores include silver, tantalum, tantalum alloys(containing Mo, Nb, Zr, W, Pt, and/or Pd), niobium, niobium alloys(containing Ta, Mo, Zr, W, and/or Pt), platinum and platinum alloys, andpalladium and palladium alloys (containing Re and Rh). Any biocompatiblematerials possessing resistivities in these ranges may be used instead.

Next, a cold drawing process may be used wherein the core 40 insertedwithin tube 42 is drawn through a die structure, resulting in a wirehaving a reduced diameter. Thereafter, the wire is annealed by heatingit to at least the beta transit temperature of the selected titaniumalloy. At this temperature, the alloy undergoes a phase transformationfrom the alpha & beta phase to full beta phase. For beta titaniumalloys, the beta transit temperature will be in a range of 600° C.-900°C. For instance, in one embodiment, Ti-15Mo has a beta transittemperature of around 730° C. Thus, annealing may occur at between 730°C.-815° C. in one example. This annealing process of using thesetemperatures to anneal the alloy changes the physical characteristics oftitanium alloy tube 42. That is, it prevents tube 42 from becomingbrittle, and will allow an additional cold-drawing step to be performedwithout the risk of the tube 42 cracking. Since the annealingtemperature of the beta titanium alloy is lower than the melting pointof the material used to form the low-resistance core (e.g., silver), thecore material will not melt when the wire is annealed. Therefore, thechallenges discussed above with respect to silver-cored MP35N wires canbe avoided.

After annealing the wire, the wire may again be cold-drawn throughanother die structure having a still-smaller diameter followed by yetanother annealing step. Any number of such iterations may be performedto obtain a wire having a desired diameter. In one example, the finalwire may have a low-resistance core that is, in one embodiment,surrounded by an unbroken layer, or tube, of the beta titanium alloy.This layer of beta titanium alloy may have a substantially uniformthickness in one embodiment. The wire may have an outer diameter ofbetween 0.001-0.01 inches. Other diameters may be used in otherexamples. Thereafter, an insulating layer may optionally be addeddepending on the application for the wire. For example, the wire may bedipped in liquefied ethylene tetrafluoroethylene (ETFE) which is thenallowed to solidify. Alternatively, such an insulating layer may beapplied using an extrusion process, such as a micro-extrusionapplication process. In a specific example, the final wire may have anouter diameter of between 0.002-0.005 inches, or about between0.0035-0.005 inches. Such an embodiment may be well-suited for some coilapplications. In another particular example, the final wire may have anouter diameter of between 0.0010 inches-0.0025 inches, which may bewell-adapted for some cable embodiments.

The process discussed above provides a wire that may have alow-resistance core and may further include a layer that contacts andsurrounds this core having a relatively low elastic modulus. In oneexample, the material forming the core may have a resistivity of lessthan 25 micro-ohm-cm (or between 10 micro-ohm-cm and 20 micro-ohm-cm ina more specific example) and the outer layer of material surrounding thecore may have an elastic modulus of between 30 GPa and 90 GPa.

Replacing MP35N with beta titanium alloy results in a wire with superiorqualities and that is better suited for medical device applications. Amedical electrical lead carrying one or more MP35N wires will besignificantly more stiff than one formed of wires made from a betatitanium alloy because of the relatively high elastic modulus for MP35Nand because of the embrittlement of the MP35N resulting from having toanneal this alloy at a temperature lower than its optimal annealingtemperature to prevent melting of the low-resistance core. The lead willtherefore be more susceptible to repeated bending and flexing, as willlikely occur in chronic implant scenarios. These factors may result inthe need to replace the lead carrying the MP35N wires more often,subjecting the patient to the inconvenience of a medical procedure.

The properties of beta titanium alloys, including their high yieldstrength, allow these materials to be readily adapted for coilapplications. As is known in the art, coils are formed by winding orgathering consecutive coil turns around a central axis. Generally, thiswinding or gathering is performed around a mandrel or other centralstructure that lies substantially along a central longitudinal axis ofthe coil. This mandrel is typically removed after the winding process iscomplete, leaving a central lumen that can be used, for instance, toreceive a stylet, guide wire, or other guiding device.

Because beta titanium alloys exhibit a high yield strength as comparedto MP35N, and further because these alloys have a very low ratio ofelastic modulus/yield strength as compared to MP35N, the beta titaniumwires are able to be used in coil configurations having a large coilpitch. As a result, the number of wires used to form the coil (that is,the number of filers in the coil) can be dramatically increased whenbeta titanium alloys are used. Moreover, for a coil having a givennumber of filars, the overall coil diameter can be decreased when a betatitanium alloy is used instead of MP35N to form the filars. Thisprovides a device (e.g., a lead) with a smaller outer diameter. Theseadvantages will be discussed further below.

Not only do beta titanium alloys result in superior coil configurations,but they also provide important benefits when cable configurations arerequired. As is known in the art, coils are formed by winding orgathering consecutive coil turns around a central axis. For instance,the winding or gathering may be performed around a mandrel in the mannerpreviously mentioned. When the mandrel is removed after the windingprocess is complete, the coil defines a tube around the central axiswhich may be described as an “air core”. If the coil is used within amedical lead, this tube may be used to receive another device such as astylet.

In contrast to coils, cables are formed by twisting multiple parallelwires together. In this case, there is no central structure around whichthe wires are twisted and there is no “air core” defined by the twistedwires after the twisting process is complete. That is, the cable is asubstantially solid structure defined by the twisted wires.

As in the case with coils, benefits exist for using beta titanium alloywires, including those containing low-resistance cores, in theproduction of cable structures. After a cable is twisted in the mannerdescribed above, the multiple wires within the cable are under stress.As a result, when the force that was applied to accomplish the twistingis removed, the wires tend to “untwist” to return to their originalparallel configuration. To allow the wires of the cable to remain“twisted together” after the cabling force is removed, the wires areheated to a stress-relieve temperature, which is the temperature atwhich the stress presented within the material is removed so that“untwisting” will not occur. Beta titanium alloys have a lowerstress-relieve temperature than MP35N, simplifying the manufacturingprocess while conserving energy. Moreover, at stress-relievetemperatures, biocompatible beta titanium alloys will not becomebrittle, resulting in ductile cables that are able to tolerate a largeamount of stress without cracking or breaking. In contrast, at itsstress-relieve temperature, MP35N does become brittle.

The embrittlement issues associated with using MP35N wires in a cableare similar to issues discussed above in relation to annealing MP35N attemperatures that are too low. In particular, embrittled MP35N wires ofa cable structure become more susceptible to repeated flexing, bending,and longitudinal force so that the lifetime of MP35N cables areshortened, possibly requiring explant of devices (e.g., leads) thatcarry such cables whereas a similar device carrying one or more betatitanium alloy cables would have a significantly longer lifetime. Theuse of beta titanium alloys in cable structures is discussed furtherbelow.

Another benefit to using beta titanium alloy wires such as thosecontaining low-resistance cores relates to their low elastic modulus (E)which ranges from 30GPa to 90 Gpa. A more specific example selects abeta titanium alloy having an elastic modulus of between 50GPa to 90Gpa. The various examples set forth above are included in one or more ofthese groups. For instance, Ti-15Mo has an elastic modulus of 80 GPa andTNTZ has a elastic modulus of 70 GPa. This is significantly lower thanthe elastic modulus for MP35N, which is 230 GPa. This results in betatitanium alloy coil and cable structures that are less susceptible tostress and strain than comparable structures made from MP35N, as shownin FIG. 5.

FIG. 5 is a diagram illustrating S-N curves for the beta titanium alloyTi-15Mo as compared to MP35N and low-titanium MP35N (LT MP35N). Duringspin-fatigue testing, Ti-15Mo and MP35N wires having diameters of 0.004inches were, during repeated cycles, subjected to strain (as a percent)shown along the y axis. The number of cycles (N) is represented alongthe x axis. The identified points along the curve indicate when failuresoccurred, although in some cases, testing was halted before the wiresfailed as indicated by the “runout” designators.

As may be appreciated from the data presented in FIG. 5, Ti-15Mo wireshave a fatigue endurance limit that is about three times higher thanthat for MP35N. For instance, point 49 of curve 50 represents a strainof a little under 0.4%. At this level of strain, an MP35N wireexperiences a failure after about 200,000 cycles. In contrast, at thislevel of strain, a Ti-15Mo wire has an infinite lifetime. Moreover,Ti-15Mo can undergo about three times the strain before a failureoccurs, as shown by curve 51 corresponding to the collection of data forthe Ti-15Mo failures. For the low-titanium MP35N wire, the results aresimilar to those of the MP34N wire, as indicated by comparing curves 50and 53.

The data shown in FIG. 5 was obtained when testing wires that did nothave low-resistance cores, but were instead wires of solid Ti-15Mo (forcurve 51), solid MP35N (for curve 50), and solid LT MP35N (for curve53). If the wires instead each included a low-resistance core of thesame diameter and composition, (e.g., all silver cores having an area of10% of the total cross-section), the curves would remain the same shapebut would each “shift-upward” slightly, with the Ti-15Mo curve 51shifting upward a little less than the other two curves 50 and 53. Whatis important to note is that regardless of whether a low-resistance coreis provided, a given Ti-15Mo wire will be able to withstand about threetimes the amount strain as its MP35N or LT MP35N counterpart.

Another benefit of using a beta titanium alloy wire relates to theability to “tune” the resistance of the wire to a desired value. In awire having an outer layer formed of MP35N surrounding a low-resistancesilver core, the resistivities of the core material and that of MP35Nare grossly mismatched. That is, the silver has a resistivity ofapproximately 1 micro-ohm-cm whereas the MP35N has a resistivity ofabout 100 micro-ohm-cm. A predetermined length of the silver-cored MP35Nwire can therefore be modeled as a first resistor having a resistance of1× in parallel with a second resistor having a resistance of 100×. Theoverall resistance of this network is largely dictated by the silvercore, with the resistance of the outer MP35N layer having very littleeffect on overall resistance.

However, if the core and the surrounding layer are formed of materialshave resistivities that are closer to one another, the resistance of thewire can be “tuned” so that it approximates a desired value. Forinstance, the resistance of a predetermined length of wire (or a cableor coil formed of one or more such wires) can be tuned to match theresistance of one or more other structures such as a connectingelectrode 36, a conducting electrode 38, or another sensor adapted tosense a signal from the body of patient (FIG. 3). As is known in thecommunication arts, matching of resistances in this manner will preventsignal reflections of the type that can obscure a sensed or atransmitted signal.

As one example of using materials that facilitate tuning, tantalum andniobium both have a resistivity of 15 micro-ohm-cm. Either one of thesematerials may be selected for use as the core while using a relativelylower-resistivity beta titanium alloy such as TNTZ having a resistivityof 90 micro-ohm-cm as the layer that directly surrounds this core. Theoverall resistance of such a wire can be tuned by adjusting the size ofthe diameter of the inner core as compared to the diameter of theoverall wire. The resistance of a given length of wire can thereby bematched to the resistance of an electrically-conducting element (e.g., aconducting or connector electrode, etc.) to which it is to be coupled,thereby improving signal transmission and reception quality. This isdiscussed further below.

Another related benefit achieved from using biocompatible beta titaniumalloy wires involves the fact that the various biocompatible betatitanium alloys exhibit a wide range of resistivities. Therefore aparticular beta titanium alloy may be selected to provide a degree ofresistivity required for a given application. Generally, it may bedesirable to have wires with resistances that are as low as possible,since this will minimize power losses and allow drive circuitry that istransmitting/receiving signals in the cable or coil structures to beoperated more efficiently.

In contrast, high-resistivity materials can be better suited for use inmagnetic resonance imaging (MRI) applications to reduce heating. Forinstance, consider a coil structure that is to be used in an MRIconditionally-safe lead. A magnetic field within the center of the coilcan induce currents that should be limited to avoid heating. Therefore,in such a scenario, it is generally desirable to form the coil of wirethat has a higher resistance to limit the induced currents flowingwithin the coil when the coil is subjected to a magnetic field. Thiscannot be readily achieved with MP35N wire having a silver core, sincethe silver has a very low resistivity. Rather, a higher resistance wirecan be provided by instead selecting a beta titanium alloy such asTi-15Mo, TLM, TNCZ or Ti-30Ta for use as the outer layer (such alloyshaving a higher resistivity than MP35N as shown in FIG. 8) with niobiumor tantalum being selected as the core, since either material exhibits asignificantly higher resistivity than silver. This can provide a coilthat is better suited for use in MRI conditionally-safe applications. Insome such applications, it may be desirable to eliminate the coreentirely to further increase resistance of the wire. This will bediscussed further below.

As is evident from the foregoing examples, in applications wherein wireresistance is an important factor in providing an acceptable solution,the wide range of resistivities that are exhibited by biocompatible betatitanium alloys makes it possible to select the right alloy to providethe desired solution. In some examples, use of a higher resistivitymaterial (relative to silver) may be used for the core to furtherenhance performance. Such materials may include, for instance, niobiumand tantalum and their respective alloys.

FIG. 6 is a graph illustrating resistivities (in ohm-cm) for variousbeta titanium alloys as well as for MP35N. As shown, MP35N has aresistivity of about 1.0×10-4 ohm-cm, or about 100 micro-ohm-cm. A widedisparity exists between this resistivity and that of silver (notshown), which is about 1 micro-ohm-cm, making a MP35N wire having asilver core difficult to tune. In contrast, the resistivity of TiOsteum,is about 9.0×10-5 ohm-cm, or 90 micro-ohm-cm and the resistivity of Tais about 15×10-5 ohm-cm, or 15 micro-ohm-cm. Because the resistivitiesof these two materials is closer to one another, a wire having a Ta coresurrounded by an outer layer, or tube, of TiOsteum can be much morereadily tuned than can an MP35N wire with a silver core.

FIG. 6 further shows the wide range of resistivities exhibited byvarious biocompatible beta titanium alloys. Thus, a beta titanium alloyshaving a desired resistivity is available for a wide variety ofapplications, such as those discussed above.

FIG. 7 is a graph illustrating tunability of a silver-cored MP35N wireas compared to that for a Ti-15Mo wire having a tantalum core. Bothwires are 10 cm (about 4 inches) in length and have a diameter of 0.004inches. The fraction of the core area relative to the core area of theentire wire (i.e., the “core area fraction”) is represented along the xaxis. The resistance of the wire (in ohms) is indicated along the yaxis. By increasing the core cross-sectional area relative to thecross-sectional area of the overall wire, the resistance of the wirewill decrease. This effect is bounded. Specifically, if the ratio ofcore cross-sectional area to overall cross-sectional area is decreasedbelow 0.05 (or about 5 percent) or increased above about 0.4 (or about40 percent), the amount of change occurring because of a change in corearea is minimal. Between the two limits, the resistance of the MP35Nwire having the silver core will range between about 2 ohms to about 0.5ohms, as shown by curve 54. For Ti-15Mo with a tantalum core, the wirecan be tuned from about 3.5 ohms to about 10 ohms, as indicated by curve55. This allows the beta titanium wire to be tuned over a significantlylarger range for a much larger variety of applications. In other examplewire configurations having outer layers formed of beta titanium alloyssurrounding low-resistance cores, a range over which tuning may occurmay generally be between 3 ohms to 15 ohms for a wire having a length ofabout 10 cm (or approximately 4 inches).

FIG. 8 is a diagram plotting resistivity against the elastic modulus (E)for various materials. The resistivity is illustrated along the y axis(in μohms-cm) and E is represented along the x axis (in GPa). As shownin this diagram, MP35N exhibits a much higher E at about 230 GPa thanany of the beta titanium alloys, as illustrated by point 56. Inparticular, Ti-15Mo has an E of 80 GPa, as represented by point 57. Thisis almost one-third the E of 230 GPa for MP35N. Other beta titaniumalloys have a similar elastic modulus. For instance, TiOsteum has anelastic modulus of about 70. The much higher elastic modulus for MP35Nindicates that MP35N is much stiffer than beta titanium alloys, andMP35N is therefore able to undergo lower fatigue for a shorter period oftime. As a result, medical devices containing wire (e.g., coil or cablestructures) made of beta titanium alloys may have a much longer lifetimethan devices carrying counterpart structures formed of MP35N. As aspecific example, high strain may occur at a site at which a lead orother device experiences repeated bending at a given location, therebycausing fatigue. A lead implanted in a patient's torso may be subjectedto such fatigue because the patient frequently bends at the waist, forinstance. This may take a toll on the life of a device (e.g., lead)having an MP35N wire, necessitating explant sooner than if acorresponding beta titanium alloy device is used.

FIG. 8 further shows that MP35N has a higher resistivity than some ofthe beta titanium alloys. In particular, MP35N has a resistivity ofabout 1.0 E-04 ohm-cm, or about 100 micro-ohm-cm, as shown by designator56. Other beta titanium alloys have a lower resistivity. For instance,Ti-30Ta-1 and TiOsteum each have a resistivity lower than that of MP35Nat about 9.0E-05 ohm-cm, or about 90 micro-ohm-cm. Ti-30Ta-1 and TNTZlikewise have a resistivity lower than that of MP35N As a result, thesebeta titanium alloys present better tuning potential than the MP35N,since their resistivity is closer to that of any low-resistivitymaterial employed for a core.

Conversely, if a high-resistivity material is desired for use in aparticular application, such as a coil configuration to be used in anMRI-conditionally safe application, a beta titanium alloy like Ti-15Momay be selected having a resistivity of about 110 micro-ohm-cm.

Other benefits relate to the biocompatibility of beta titanium alloyssuch as Ti-15Mo as compared to MP35N. This is shown in Table 2, whichdescribes the metal ion release of MP35N and Ti-15Mo when submerged insaline for four weeks.

TABLE 2 Ion Release (μg/mm²) MP35N Ti—15Mo Ti — 0 Mo 0.00013 0.00018 Co0.0022 — Ni 0.0022 — Cr 0.0000 —

As shown in Table 2, when submerged in saline solution, the total metalion release occurring from MP35N is more than twenty-five times higherthan for Ti-15Mo over a four-week period. The highest contributors tothis ion release are Co and Ni, with release levels being 0.0022 μg/mm.²In general, this level of ion release is not exhibited by biocompatiblebeta titanium alloys. In particular, since biocompatible beta titaniumalloys such as Ti-15Mo do not contain Co and Ni, the overall ion releaseis much more limited.

Furthermore, because Ti-15Mo and other beta titanium alloys do notcontain Co, such alloys do not exhibit metal ion oxidation (MIO). MIOoccurs when the Co contained by MP35N ionizes and infuses into thesurrounding insulating layer 46 (e.g., polyurethane). This reduces thelifetime of the insulating layer, making it necessary to explant thedevice (e.g., the lead) sooner than would otherwise be required for adevice that instead carries wires made of biocompatible beta titaniumalloys such as Ti-15Mo which do not contain Co and thus do not exhibitMIO.

FIG. 9 is a graph of the level of biocompatibility for various elementsas compared to polarization resistance (N. Niionmi, Journal of Metal,June 1999, pp. 32-34). Biocompatibility is illustrated along the x axis.Polarization resistance (R/Ω-cm), which is generally inverselyproportional to the corrosion rate, is shown along the y axis. Elementssuch as Ti, Mo, Nb, Ta, Zr, Cr, Fe, and Sn included in biocompatiblebeta titanium alloys exhibit superior biocompatibility as compared toelements Co and Ni of MP35N. This is particularly important whereapplications involving leads to be implanted within the brain areconcerned. For instance, leads configured to deliver deep brainstimulation (DBS) therapy optimally are associated with a minimum levelof ion release such as is exhibited by biocompatible beta titaniumalloys of the type described herein.

Still another benefit of using beta titanium alloys rather than MP35Nrelates to price. This is shown by comparing costs of Ti-15Mo to MP35N.While the two are substantially the same price per pound, Ti-15Mo ismuch less dense than MP35N. Specifically, a cubic centimeter of Ti-15Moweighs 4.95 grams versus 8.43 grams for the same volume of MP35N. As aresult, it is estimated that for about a same amount of money, Ti-15Mowill produce about 70% more wire than MP35N.

Additionally, beta titanium alloys are weld-compatible with variousmaterials typically included in implantable devices such as Nb, Nballoys, Pt, Pt alloys, Ta and Ta alloys. This allows bonds to be readilycreated with such materials. In contrast, MP35N is not weld-compatiblewith these materials. As a result, it may take longer to form a goodweld between a wire formed of MP35N and one of these materials.Additionally, more MP35N pieces may fail inspection.

Some of the various characteristics of Ti-15Mo and MP35N described aboveare summarized in the following Table 3:

TABLE 3 Characteristic MP35N Ti—15Mo Metal Ion >.0045 micrograms/ .00018micrograms/ Release mm² mm² Biocompatibility Ion Release of about 25XIon release of about X Manufacturability Weld-incompatible withWeld-compatible with Nb, Nb alloys, Pt, Pt Nb, Nb alloys, Pt, Pt alloys,Ta and Ta alloys alloys, Ta and Ta alloys Density 8.43 grams/cm³ 4.95grams/cm³ Length of Wire L Approximately 1.7L for a given cost

The foregoing, and other benefits to be discussed below, are provided byusing a biocompatible beta titanium alloy having an elastic modulus ofbetween 30 GPa and 90 GPa to form wires to be used within medicaldevices including, but not limited to, leads or lead extensions. In someexamples, these wires optionally have low-resistance cores. Such wiresmay be used in a variety of configurations, including coilconfigurations or cable arrangements, as will be discussed furtherbelow. For example, in FIG. 3, a single such wire may extend betweeneach connector electrode 36 and a corresponding conductor electrode 38such that the body of the device carries four wires, each beingelectrically insulated from the other wires so that independent signalsmay be carried by each wire. Alternatively, if redundancy is required,more than one wire may extend between each such connector and conductorelectrode pair, with such wires being electrically coupled one toanother. In this latter case, more than four wires may be carried by thelead body of FIG. 4 to provide connectivity between the four connectorand conducting electrode pairs 36, 38. Within the lead body, themultiple wires may be twisted together in a cable arrangement or mayinstead comprise a coil. The various embodiments are discussed furtherbelow.

Returning now to FIG. 4B, this figure illustrates a side cutaway view ofa distal end of a lead 47 that carries a beta titanium wire having alow-resistance core 40 as described herein electrically and mechanicallycoupled to an electrode 43. Lead 47 is shown to carry a single wire 39having a low-resistance core 40 surrounded by tube 42 which may beformed according to methods described herein. To affix an end of wire 39to an inner surface of electrode 43 (also shown cut-away), the end ofthe wire may be abutted to the inner surface of the electrode 43 and thetube 42 may be heated to create an electrical bond with the innerelectrode surface. In one embodiment, one or more conductive joints 45may be formed on one or more sides of tube 42 to facilitate thisbonding, as by melting one or more beads of material having a highconductivity.

Within the lead body, wire 39 may extend substantially the length oflead body to electrically and mechanically couple to a connectorelectrode 36, such electrode being adapted to interface with a connectorof an implantable device such as IMD 12.

Lead 47 may carry only a single conducting electrode 43 that is coupledto a corresponding connector electrode (not shown in FIG. 4B) via wire39. Alternatively, lead may carry additional conducting electrodes, eachin a manner similar to that shown in FIG. 3, with each conductingelectrode being coupled to a corresponding connector electrode 36 via anassociated wires in some embodiments. Each such wire may have a coreformed of a low-resistivity material encased by a layer of beta titaniumalloy, as previously disclosed. As is possible with the beta titaniumalloy, each such wire may have a resistance that is tuned to acorresponding structure to which it is connected. For instance, in FIG.4B, the resistance of wire 39 may be tuned to that of electrode 43. In amore specific example, electrode 43 has a same, or similar resistance asa corresponding connector electrode located at a proximal end of leadsuch that the resistance of wire 39 is tuned to both of theseinterconnecting structures.

As discussed above, multiple wires of the type described herein may beconfigured in either a coil or a cable arrangement. When configured ineither configuration, the beta titanium alloy wires provide significantbenefits as compared to low-resistance core wires that are insteadformed of MP35N, as has traditionally been the case. The followingdiscussion will therefore specifically consider examples of betatitanium alloy wires arranged in coil configurations. Thereafter, cableconfigurations will be described.

FIG. 10A is a side view of one embodiment of a coil 69 that is formed ofat least one filar 60 (or wire) of the type described herein having abeta titanium outer layer and a low-resistance core. For purposes ofFIG. 10A, it will be assumed one wire is being wound to form thesingle-filar coil, although more filars may be used as discussed inreference to FIG. 10B or 10C.

In one embodiment, it is desirable to have a filar that is as thin aspossible for use in coil 69 since a coil formed of a small diameterfilar will withstand more bend stress than one formed of the samematerial but which has a larger diameter.

A filar suitable for coil 69 can be obtained by an iterative processthat draws the wire comprising beta titanium tube 42 and alow-resistance core 40 (FIG. 4A) through a die, followed by an annealingprocess to anneal the beta titanium tube. In one example, this annealingwill occur at between 600-900° C. for beta titanium alloys. Inparticular, for Ti-15Mo, the annealing temperature is about 730° C. Atthese temperatures, the beta titanium will be completely annealed andwill not become brittle. This process may be repeated multiple times toobtain a filar of the desired diameter.

The wire may have a starting inner diameter ranging from 0.1 inches to 2inches (or between about 0.25 cm-5.0 cm). A low-resistance core may beinserted into the tube, with the core being formed of a material havinga resistivity of less than 25 micro-ohm-cm. Following the iterativedrawing and annealing steps, the diameter of the wire may be betweenabout 0.001 inches and 0.01 inches. In a more specific example, theending diameter of the filar may be between about 0.001 inches and 0.002inches. In a still more specific example, this diameter is about 0.0013inches.

As discussed above, one advantage with using the biocompatible betatitanium alloy wire as opposed to an MP35N wire is that MP35N must beannealed at temperatures above 1000° C. to prevent brittleness. However,at this temperature, the low-resistance core 40 will melt, causingmanufacturing challenges. This is not an issue for a low-resistance corebeta titanium alloy wire, since this wire may be fully annealed belowthe melting temperature of the core and the resulting wire will not bebrittle following the annealing process.

As shown in FIG. 10A, after being drawn to a desired diameter, filar 60may be wound to form a coil. As defined herein, and as known in the art,a “coil” refers to a series of multiple connected turns formed about acentral axis 67 by gathering or winding. This is opposed to a cablestructure that is formed by twisting multiple substantially parallelwires together to form a rope-like configuration.

During a coil winding process, filar 60 may be wound around a mandrel(not shown in FIG. 10A) which is removed after the winding process iscomplete. The filar 60 is wound at a desired pitch angle θ63 that may bemeasured with respect to a cross-sectional axis 65 that is perpendicularto the longitudinal axis 67 of coil 69.

When the pitch angle θ63 is selected to be zero such that the coilwindings are substantially parallel to cross-sectional axis 65,conductor coil 69 is able to withstand a high level of stress. In thiscase, when force is exerted on filar 60, conductor coil 69 is readilyable to expand a maximum amount without breaking or becoming permanentlydeformed.

However, disadvantages exist to having a smaller pitch angle. Forinstance, the amount of material required to form coil 69 increases aspitch angle θ63 decreases. As a result, the total weight and overallresistance will increase with the decreasing pitch angle. The increasingresistance will, in turn, result in higher power losses over the lengthof the coil. Moreover, outer diameter will increase as the pitch angleθ63 decreases.

While selecting a pitch angle θ63 that is somewhat above zero willdecrease material costs and weight of the product while reducing powerlosses and outer diameter, selecting a pitch angle that is too largewill increase manufacturing complexity. This is particularly true whencoil 69 is formed of a single filar 60 as shown in FIG. 10A (rather thanof multiple filars as will be described in reference to FIG. 10B). Whencoil 69 is formed of a single filar, it is difficult to maintain thesingle filar at an angle that is above about 30° during the windingprocess. To achieve a greater pitch angle in a single-filar embodiment,a special winding mechanism must be used that increases productioncosts. Therefore, attempting to increase the pitch angle too much tosave material costs will have the adverse effect of increasingmanufacturing complexity. For this reason, when coil 69 is formed of asingle filament, it is desired in some examples to select pitch angleθ63 to be between zero and 30 degrees.

As mentioned above, in some embodiments, more than one conductorfilament is used to form coil 69. This may be desirable, for example, ina lead having multiple conducting electrodes. In such a case, a coilhaving multiple filars that are each insulated one from another may beemployed to electrically couple the multiple conducting/connectorelectrode pairs of the lead. That is, each conducting electrode 38 maybe coupled to a respective filar that carries an outer insulating layer46, as discussed in reference to FIG. 4A. The multiple filars may bewound into a single coil structure, as described in reference to FIG.10B.

FIG. 10B is a side view of a coil 71 according to another embodiment ofthe current disclosure. In this example, coil 71 is being wound oftwelve filars 66 a-66 l, although more or fewer filars may be used inthe alternative. In one embodiment, each of the twelve filars hassubstantially the same outer diameter, which may be between about 0.001inches-0.01 inches. In a more specific example, the wires are betweenabout 0.0035 inches-0.005 inches. Each filar may be formed by colddrawing, then annealing, a beta titanium tube surrounding alow-resistance core in a manner similar to that described above. Whilenot shown in FIG. 10B, in one embodiment, each of filars 66 a-66 lincludes an insulating layer 46 (FIG. 4A) so that each filar isinsulated from adjacent filars and may independently transmit andreceive signals.

The filars may be positioned side-by-side and coiled at a same time, asshown in FIG. 10B. Winding all filars at once in this manner ispreferable because it saves time and helps maintain alignment. While thefilars may instead be wound one at a time, the winding process wouldtake longer to complete, and the spacing of the filars would be moredifficult to control, adding complexity to the process.

In this example, the filars are being wound around a mandrel 62 thatprovides structure to maintain the shape of coil 71 during the windingprocess. Mandrel 62 may be removed upon completion of the winding,leaving an “air-core” coil. The center of the coil defines a space thatmay, for instance, receive a stylet or guide wire to aid in placing adevice (e.g., a lead) that carries coil 71 at a desired location withina patient's body.

As is the case with the embodiment of FIG. 10A, filars 66 a-66 l may bewound at a pitch angle θ63 measured from cross-sectional axis 65. Ingeneral, as the number of filars increases, both the pitch angle θ63 andthe pitch 64 must increase, wherein the pitch is the distance betweenadjacent turns of the same filar. Both the pitch angle θ63 and the pitch64 is limited by the flexibility of the filar. Wires made of materialshaving a higher elastic modulus such as MP35N cannot accommodate aslarge of a pitch angle or pitch as those formed of beta titanium alloyshaving a much lower elastic modulus. Thus, coils of a predetermineddiameter that are formed of MP35N wire cannot contain as many filars asthose of a same diameter that are formed of beta titanium wire.Therefore, coils of a predetermined diameter that are formed of MP35Ncannot provide as many electrical connections and are able toaccommodate fewer electrodes when compared to a similar coil formed ofbeta titanium alloys.

The foregoing may be stated another way: a coil formed of the stifferMP35N wires having a predetermined number of filars will require alarger diameter than a coil formed of beta titanium alloy having thesame number of wires. For instance, if coil 71 is formed of eight MP35Nwires, an outer diameter that is greater than 0.032 inches must beobtained if the pitch is to be limited to that which will not placeundue strain on the wires. In contrast, if coil 71 is instead formed ofeight beta titanium alloy wires, a coil having an outer diameter ofabout 0.016 inches, half that of the MP35N coil, can be obtained.Minimizing the size of a coil may allow the size of an associated device(e.g., a lead carrying the coil) to be substantially reduced, somethingthat is very important for medical device applications. Theseconsiderations are discussed further in regards to FIG. 11A below.

Still another benefit of using a beta titanium wire to form coil 69, 71involves the bend radius at yield that can be achieved by this coil ascompared to a coil of similar dimensions formed of MP35N wire. The bendradius at yield measures the minimum radius that can be achieved by thecoil without kinking or damaging the coil. Because a beta titanium alloywire used to form the coil has a much lower elastic modulus, theresulting coil structure is not as stiff and has a smaller bend radiusat yield than a similar MP35N coil. Therefore, a medical device such asa lead, lead extension, stylet, guide wire, or any other type of devicethat carries such a coil will likewise not be as stiff, and will be ableto much more readily navigate the twists and turns of the human anatomy.A comparison of the bend radius at yield of an MP35N coil as opposed tothat for a beta titanium alloy coil is provided below in reference toFIG. 11A.

FIG. 10C is a conceptual drawing of a side cut-away view of a coilcontaining eight filars 80 a-80 h (“filars 80”), each coupled to arespective one of ring electrodes 82 a-82 h (“electrodes 82”, shown incross-section in FIG. 10C), wherein the electrodes are carried by adistal end of a medical electrical lead 78. For instance, filar 80 a isshown coupled to electrode 82 a, filar 80 b is shown coupled toelectrode 82 b, and so on. Each of the filars may be formed of a betatitanium alloy having a low-resistance core as described herein. Therelatively large number of filars carried by the coil is made possibleby the relatively large pitch and large pitch angle θ that is affordedby these filars, something not achievable when using conventional MP35Nfilars.

In one example, each of filars 80 a-80 h is provided with an insulatinglayer 46 (FIG. 4A) that insulates it from adjacent filars, therebyallowing each filar to carry an independent electrical signal. Asdiscussed above, the insulating layer may be a polymer jacket or a layerformed of some other insulating material. Moreover, lead 78 may includean insulating lead body 83. This may be formed of any of thebiocompatible polymer materials suitable for medical electrical leads,as described herein.

According to one example, lead 78 has a proximal end (not shown)carrying eight connector electrodes each similar to electrodes 36 ofFIG. 3, and each being coupled via a respective one of filars 80 a-80 hto a corresponding one of conducting electrodes 82 a-82 h. Theseconducting electrodes may be electrically and mechanically coupled via arespective connecting electrode to corresponding connectors within aheader of IMD 12. The IMD may thereby transmit signals to a patient todeliver therapy and/or to receive signals sensed from the patient formonitoring and diagnostic purposes, for example.

IMD 12 may contain a stimulation pulse generator that may be capable ofdelivering stimulation simultaneously via eight separate stimulationchannels so that each of conducting electrodes 82 a-82 h provides anindependent stimulation signal to patient 10 at a same time. In anotherscenario, IMD 12 may drive electrodes 82 a-82 h in a time-multiplexedmanner such that not all of electrodes 82 a-82 h are being drivensimultaneously. In still another example, IMD 12 may drive only aselected subset of electrodes 82 a-82 h, that subset being selectedbased on patient response or some other indication of a level ofefficacy of therapy, level of paresthesia, and/or level of side effectsresulting from stimulation via that electrode combination.Alternatively, multiple electrodes may be driven with a same signal. Inyet another example, some of the electrodes 82 may be used to deliversignals to tissue of patient 10 while other electrodes 82 may insteadsense signals. As yet another example, all or a subset of all, ofelectrodes 82 may be used to deliver stimulation some of the time whilesensing signals at other times in a time-multiplexed manner. Suchcontrol options may depend on the stimulation and sensing capabilitiesprovided by an IMD 12, which may be therapy dependent. In any event,lead 78 carrying the improved coil comprising filars 80 providesenhanced capabilities to deliver and/or sense more signals to/from morelocations.

As previously described, filars 80 may each have a resistance that istuned to that of electrodes 82 in the manner described above. In oneexample, not all electrodes 82 need have the same resistance, and thefilars may have a resistance tuned to the electrodes to which they arecoupled such that not all filars have a same resistance in one example.

It will be understood that FIG. 10C is a conceptual drawing that is notnecessarily to scale. For instance, the ring electrodes may besubstantially wider than a diameter of a single filar, as shown. What isimportant to note is that one or more filars may each be electricallyand mechanically coupled to a respective electrode. This may beaccomplished by removing a portion of an insulating layer that surroundseach of the filars to expose a portion of that filar (e.g., exposing anend of filar 80 a). In one example, the exposed end of the filar may beabutted next to an inner surface of a respective conducting electrode(e.g., electrode 82 a). This exposed portion of the filar may be heatedby a resistive-, spot-, or a laser-welding process. This will bond theexposed portion of the filar to this surface of the electrode, formingan electrical and mechanical connection. This bonding may be furtherstrengthened by melting beads of a fusible metal alloy to formelectrically conductive joints 86. For instance, beads of ahigh-conductivity material may be melted to form these conductivejoints. As previously noted, this can be accomplished more readily withthe beta titanium wire because this wire is weld-compatible with variousmaterials typically included in implantable devices such as Nb, Nballoys, Pt, Pt alloys, Ta and Ta alloys. Similar bonds may be createdwith other interconnecting structures such as connector electrodesand/or sensors.

In other embodiments, the bond between a filar and a correspondinginterconnecting structure (e.g., an electrode or sensor) may be createdin another way. For instance, a mechanical coupling mechanism may beused instead of, or in addition to, a weld process. Such a mechanicalcoupling may involve crimping, pinching, threading, tying, or otherwisemechanically affixing a portion of the wire to the interconnectingstructure to form a mechanical and electrical connection.

FIG. 10C shows each of electrodes 82 being coupled to a differentrespective one of filars 80. This may be desirable in an embodimentwherein a large number of such conducting electrodes are being coupledto connecting electrodes at a proximal end of the lead. In such a case,each of filars may be electrically insulated from the other filars toallow the conducting electrodes to be transmitting signals independentlyof the other electrodes, if desired. In this case, redundancy may not bepossible since all of the available wires are needed to accommodate themultiple conducting electrodes. In another example, two or more filarsmay be coupled to a same electrode 82 to provide redundancy.

The lead of FIG. 10C may take many embodiments. For instance, electrodes82 need not be ring electrodes but could be segmented electrodes that donot extend around the entire circumference of lead 78 such as shown inFIG. 3B and that optionally have a complex array geometry. Examples ofcomplex array geometries are described in commonly-assigned applicationSer. No. 11/591,188 filed Oct. 31, 2006 entitled “Programming Interfacewith a Cross-Sectional View of a Stimulation Lead with Complex ElectrodeArray Geometry”, the entire content of which is incorporated herein byreference. Such geometries include electrodes occupying a samelongitudinal position on a lead (such as shown in cross-section in FIG.3B) as well as electrodes occupying different longitudinal positions onthe lead. Additionally or alternatively, one or more sensors may becarried by lead 78, each being coupled to one or more of filars 80. Suchsensors could include any of the sensors known in the art, such aspressure transducers, temperature sensors, sensors for measuring a levelof a substance within the patient's body, motion and/or activitysensors, and so on.

As previously described, during the manufacturing process, filars 80 maybe coiled around a mandrel 62 (FIG. 10B) which is removed after windingis completed, with the inner coil diameter defining an inner lumen. Inan alternative example, filars 80 may be coiled around a structure thatdefines an inner lumen and which is not removed after winding. In eithercase, such a lumen may accommodate a stylet, guide wire, or anothersteering device that may be used to position the coil at a desiredlocation within a patient's body.

While FIG. 10C illustrates eight conducting electrodes 82, each beingcoupled to a different respective one of the multiple filars 80 of thecoil 78, it will be appreciated that a similar coil having a differentnumber of filars may be configured in a similar manner. For instance,the twelve-filar coil 71 of FIG. 10B may be used to couple twelveconducting electrodes each to a respectively different connectorelectrode if desired so that twelve signals may be transmitted and/orreceived simultaneously (assuming that capability is provided by IMD12). In another embodiment, even more than twelve filars may beprovided.

The following Table 4 compares characteristics of an eight-filar coilformed of MP35N as compared to an eight-filar coil formed of a betatitanium alloy Ti-15Mo.

TABLE 4 Characteristics for Eight-Filar Coils MP35N eight- Ti—15Moeight- Characteristic filar coil filar coil Fatigue life at strain of.38% 200,000 cycles Infinite Pitch (with same OD) X 2X Pitch AngleLimited range Wider range due to high stress tolerance Coil OuterDiameter At least 2D D

Table 4 shows that the fatigue life of a MP35N coil (assuming filarsthat are 0.004 inches in diameter) at a strain of about 0.38% is 200,000cycles. In contrast, at this same strain, the Ti-15Mo coil has aninfinite fatigue life. As such, an MP35N coil experiencing this strainwill need to be replaced after a predetermined period of time, whereasthis will not be necessary for a similar Ti-15Mo coil. Leads or othermedical devices therefore have a longer implant life when a Ti-15Mo coilis used.

The fatigue data of Table 4 represents coils formed of filars withoutlow-resistance cores. However, a similar comparison may be drawn betweenMP35N coils and Ti-15Mo coils if each are formed of filars havinglow-resistance cores. In particular, an MP35N coil formed of filarshaving low-resistance cores may tolerate somewhat more strain at 200,000cycles than an MP35N coil formed of filars without a core. However, thisadditional strain tolerance will likewise be manifested in the Ti-15Mocoil formed of filars having low-resistance cores so that the differencein performance between the two coils remains the same.

Table 4 further compares the pitch of the two coils assuming an outercoil diameter that is the same. As discussed above in reference to FIG.10B, the pitch describes the distance that may be obtained betweenadjacent turns of a same wire within a coil structure, as indicated byarrow 64. Pitch is affected by the strain tolerance of the material, asillustrated in FIG. 5. Beta titanium alloys are much more straintolerant than MP35N. In particular, beta titanium alloys have a higherendurance limit such that they may be subjected to a relatively largeamount of strain without permanently deforming or breaking. This isfurther quantified by their very low ratio of elastic modulus/yieldstrength, which is a characteristic required for forming multiple-filarhigh-pitch coils. Because of these characteristics, the pitch of thebeta titanium alloy coils can be significantly increased over theirMP35N coil counterparts. For instance, as shown in Table 4, the maximumpitch of an eight-filar MP35N coil having a given diameter will be abouthalf the pitch that can be achieved for the eight-filar Ti-15Mo coil.This increased pitch allows the Ti-15Mo coil to include more filars thanan MP35N coil having the same diameter.

Next, pitch angle is considered for both the MP35N coil and its Ti-15Mocounterpart. As previously discussed, the pitch angle θ63 (FIG. 10B)describes the angle between the coil filars and an axis perpendicular tothe longitudinal axis of this coil, also referred to as the“cross-sectional axis”. As shown, an MP35N coil has a limited pitchangle range. The pitch range is increased for Ti-15Mo coils, since theyhave a higher stress tolerance. As the pitch angle θ63 increases, morefilars may be accommodated by a coil of a predetermined diameter. Statedanother way, an MP35N coil having a same number of filars as a Ti-15Mocoil will have a much larger diameter than the Ti-15Mo coil. This istrue, at least in part, because the pitch angle for the MP35N coil mustbe smaller than that for a Ti-15Mo coil. For instance, as previouslydiscussed, while the outer diameter of an MP35N coil having eight filarsis greater than 0.032 inches, the outer diameter of the Ti-15Mo coilthat also comprises eight filars can be reduced to 0.016 inches. This isa significant size reduction which can provide important benefits formedical electrical leads and other devices that optimally have as smalla diameter as possible to aid in device placement, increased patientcomfort, and enhanced cosmetic appeal.

In addition to the benefits shown in Table 4, other benefits exist forusing beta titanium alloy filars in coil designs. As previouslydiscussed above, Ti-15Mo coils are easier to manufacture than the MP35Ncounterparts for several reasons. First, in the embodiment wherein theTi-15Mo wires have a low-resistance core, a full annealing process canbe performed for the Ti-15Mo filars since the annealing temperature islower than the melting point of the materials which may be selected forthe lower resistivity core. Because the Ti-15Mo wire can be fullyannealed without risk of melting the core, the resulting coil structurewill not be brittle. This is not the case for MP35N coils formed offilars having a low-resistance core. Because the MP35N filars must beannealed at a lower-than-desired annealing temperature to preventmelting of the core, the resulting filar (and hence the coil) is brittleand cannot withstand the same amount of fatigue as its biocompatiblebeta titanium alloy counterpart. Thus, coils that include MP35N filarsare not as flexible and do not withstand the same amount of fatigue asthe beta titanium alloy coils.

FIG. 11A is a graph comparing the bend radius at yield of two coils(described in inches along the x axis) to the maximum pitch of the coils(represented in inches along the y axis.) In this example, each of thetwo coils has an outer diameter of 0.027 inches, with the filars of eachof the coils having a diameter of 0.004 inches. Point 90 represents acoil formed of one or more Ti-15Mo filars having a maximum pitch of 0.08inches when the bend radius at yield is about 0.067 inches. At thispitch, up to twelve filars can be included in the coil. Each such filarmay be individually insulated such that up to twelveconducting/connector electrode pairs can be individually andindependently interconnected using this coil. That is, twelveindependent signals may be transmitted and/or received at once usingthis coil. In contrast, point 92 represents a coil formed of up to fourMP35N filars. The filars of this coil have, at most, a pitch of about0.027 inches when the bend radius at yield is about 0.0167 inches. Atthis pitch, only four filars, at most, can be accommodated. In otherwords, assuming each filar is individually insulated one from another,this coil can carry, at most, four independent signals, which wouldsupport only four conducting/connector electrode pairs.

While FIG. 11A applies specifically to a coil having an outer diameterof 0.027 inches, the results of FIG. 11A can be generalized in somerespects. This generalization can be drawn by determining the ratio ofmaximum pitch to the coil outer diameter. For MP35N, the maximum pitchis about 0.027 inches, as indicated by point 92, resulting in a ratio ofmaximum pitch to outer diameter of 1. For MP35N, when this ratio exceeds1, the strain increases significantly, resulting in a structure thatwill have a significantly shorter lifespan.

A similar ratio of maximum pitch to coil outer diameter may beapproximated for Ti-15Mo. In particular, the ratio of maximum pitch toouter diameter is 2.96 (very close to three). Thus, the Ti-15Mo coil caninclude significantly more filars (2 to 3 times as many) for a givencoil diameter, or conversely may have a much smaller diameter (2 to 3times smaller) for the same number of filars as compared to its MP35Ncoil counterpart. These performance benefits can be provided for up tothe number of filars (twelve in this example) represented by point 90 ofFIG. 11A before the strain will increase significantly for the Ti-15Mocoil.

In implantable medical device applications, providing a device (e.g., alead) that has a diameter that is as small as possible while providingfor transmission of as many signals as possible is highly desirable.Moreover, it is beneficial to provide these characteristics without anincreased strain which will shorten the lifespan of a device. As can beseen by the data of FIG. 11A, beta titanium alloys such as Ti-15Mo coilscan provide for both of these objectives, supporting both a large filarcount while minimizing outer diameter up to the illustrated limits andproviding a smaller bend radius at yield. These objectives cannot be asreadily achieved with MP35N coils.

Next, the bend radius at yield for MP35N coils and Ti-15Mo coils can beconsidered in more detail. As shown with respect to point 90, at a pitchof 0.08 inches, a Ti-15Mo coil having up to twelve filars can achieve abend radius of about 0.067 inches at yield (that is, before permanentdeformation of the coil occurs.) Thus, even when the pitch is high, avery small bend radius can be achieved by a coil comprising up to twelveTi-15Mo filars. In contrast, point 92 represents a coil formed of only,at most, four MP35N filars, although the coil could contain fewerfilars. At yield, the MP35N coil can only achieve a bend radius of 0.167inches when the filars are coiled at a pitch of about 0.027 inches.Thus, the beta titanium coil has a bend radius at yield that is about2.5 times smaller than that of the MP35N coil of the same 0.027 inchouter diameter. This is true even though the beta titanium coil has upto three times the number of filars. Because of the low modulus of thebeta titanium alloy, the Ti-15Mo coil is less stiff and can undergo moreflexing without permanent deformation than can the MP35N coil. Theoverall result is that the Ti-15Mo coil is smaller and more flexible,allowing it to more readily navigate through the tortuous paths of thehuman body during device placement (e.g., when a lead is being navigatedinto position in association with a therapy target.)

Point 90 represents the maximum pitch and minimum bend radius at yieldthat can be achieved by a Ti-15Mo coil having an outer diameter of 0.027inches. It may be noted that a Ti-15Mo coil having this diameter cansupport any bend radius down to 0.067 inches and any pitch up 0.08inches for a coil having up to twelve filars. Moreover, any number offilars fewer than twelve may be included in a Ti-15Mo coil having thisouter diameter and such a coil will likewise be able to achieve a bendradius of 0.067 inches with a pitch of 0.08 inches. As discussed above,each filar of such a coil may be individually insulated from otherfilars to carry an independent signal.

In one example, a four-filar coil formed of a beta titanium alloy mayhave a pitch of up to 0.08 inches with an outer diameter of 0.027inches. At this pitch and with only four filars, adjacent turns of thefilars will not be touching one another in the manner shown in FIG. 10B.Instead, “gaps” will exist between a turn of one filar and an adjacentturn of another filar. One benefit of using this type of high-pitchconfiguration wherein adjacent filars are not tightly packed involvesthe overall length of each filar. In particular, each filar will have atotal length that is shorter than it would otherwise be if wound at adecreased pitch and with adjacent turns of a filar contacting oneanother. This decreased length of the filar in turn decreases theoverall resistance of the filar, which is advantageous for many medicalapplications. Moreover, the decreased filar length further reduces thematerial costs needed to produce the filar. Additionally, if adjacentfilars are not tightly packed such that adjacent filar turns do notcontact one another when the coil is in a substantially “straight”configuration, certain coil configurations may have a larger bend radiusat yield. This is because physical contact between the adjacent filarswill not be a limiting factor in preventing a coil from flexing.

The data shown in FIG. 11A is provided for coils having filars that aresolid MP35N and solid Ti-15Mo. However, similar results will be achievedfor coils having filars with cores formed of a low-resistivity material.That is, if cores formed of a low-resistivity material of a same sizeare included in an MP35N coil as well as a Ti-15Mo coil, the relativecomparison between such coils will remain the same as that shown in FIG.11A. Specifically, the Ti-15Mo coil of an outer diameter of about 0.027inches will still be capable of providing for three times the pitch asthe MP35N coil of the same outer diameter. Moreover, the Ti-15Mo coilwill still achieve a bend radius that is about 2.5 times smaller thanthe MP35N coil when both coils have an outer diameter of about 0.027inches.

The data of FIG. 11A is specifically related to Ti-15Mo coils having amaximum pitch for the listed bend radius at yield. Similar results willbe obtained for other beta titanium alloys. In other examples, betatitanium alloy coils having similar dimensions as those shown in FIG.11A will exhibit up to a 0.1 inch maximum pitch while having a similarbend radius at yield. Thus, the superior results obtained for Ti-15Mocoils are likewise obtained for coils formed of other beta titaniumalloys.

FIG. 11B is a graph comparing the bending radius at yield for afour-filar MP35N coil to that of a four-filar Ti-15Mo coil. The filarsin this case have a diameter of 0.004 inches and the coils have a pitchof 0.03 inches. The bending radius at yield (in inches) is shown alongthe Y axis. The coil outer diameter (in inches) is depicted along the Xaxis. Curve 94 represents data obtained for the Ti-15Mo coil whereascurve 95 represents data obtained for the MP35N coil.

As shown by curve 95, the MP35N coil has a bend radius of about 2.5times that of a corresponding Ti-15Mo coil having a same coil diameter.For instance, for a coil outer diameter of 0.027 inches, the Ti15Mocoil's bending radius at yield is about 0.043 inches whereas the coilbending radius at yield for a MP35N coil having the same outer diameterand pitch is about 0.11 inches (which is about 2.5 times that of theTi15Mo coil). Similarly, at a coil outer diameter of 0.05 inches, theTi15Mo coil's bending radius at yield is about 0.02 inches whereas thecoil bending radius at yield for a MP35N coil having the same outerdiameter and pitch is about 0.05 inches (again, about 2.5 times that ofthe Ti15Mo coil). This again shows the superior ability of the Ti-15Mocoil to bend without permanent deformation as compared to the MP35Ncoil. This is important in applications wherein the coil will besubjected to repeated strain with small bend radius.

FIG. 11C is a graph comparing the bending radius at yield for aneight-filar MP35N coil to that of an eight-filar Ti15Mo coil. The filarsin this case have an outer diameter of 0.004 inches and the coils have apitch of 0.07 inches. The bending radius at yield (in inches) is shownalong the Y axis. The coil outer diameter (in inches) is depicted alongthe X axis. Curve 96 represents data obtained for the Ti-15Mo coilwhereas curve 97 represents data obtained for the MP35N coil.

As was the case with the four-filar coils represented by FIG. 11B, theMP35N coil has a bend radius of about 2.5 of a corresponding Ti-15Mocoil having a same coil diameter. For instance, for a coil outerdiameter of about 0.039 inches, the Ti-15Mo coil's bending radius atyield is about 0.05 inches whereas the coil bending radius at yield fora MP35N coil having the same outer diameter and pitch is about 0.125inches (which is about 2.5 times that of the Ti15Mo coil). Similarly, ata coil outer diameter of 0.05 inches, the Ti15Mo coil's bending radiusat yield is about 0.04 inches whereas the coil bending radius at yieldfor a MP35N coil having the same outer diameter and pitch is about 0.10inches (again, about 2.5 times that of the Ti-15Mo coil).

Other advantages to using beta titanium alloys as compared to MP35N forcoil applications are exemplified by FIG. 12.

FIG. 12 is a graph comparing the coil pitch to the total length of wirerequired to form a coil that is 40 inches in length having an outerdiameter of 0.027 inches. As pitch increases (as shown along the xaxis), the total length of material that is required to form the coildecreases (as indicated along the y axis). For instance, at a pitch of0.027 inches, a little over 120 inches of wire is required to form thecoil, whereas at a pitch of 0.039 inches, a little under 90 inches isrequired. Thus, over 25% less material is required for the coil havingthe larger pitch. This illustrates the advantage of using a materialsuch as a beta titanium alloy that is capable of providing a wire havinga very large pitch.

Besides decreasing material costs, the reduction in length achieved byincreased pitch is important for another reason. The 25% reduction inoverall length of a filar of a coil attained by increased pitch willdirectly translate into a 25% reduction in the resistance of the filar,helping to minimize power losses and overall device resistance.

Still another advantage of using Ti-15Mo coils relates to the elasticrange provided by these coils, which is almost twice that of coilshaving MP35N filars. This is shown in FIG. 13.

FIG. 13 is a graph illustrating an enlarged strain plot for a Ti-15Mocoil and an MP35N coil, wherein both coils have the same outer diameter.An amount of energy that is expended to exert strain on a coil is shownalong the y axis in lbf-in. The amount of strain the coil undergoes (asa percentage) is represented along the x axis. Curve 100 illustratesthat the Ti-15Mo coil has a much larger elastic range than that for theMP35N coil represented by curve 101. In particular, at an energyexpenditure of 0.0002 lbf-in, the MP35N coil experiences a strain of7.5%, versus a strain of about 15% for Ti-15Mo. Similarly, when 0.0004lbf-in is expended, the MP35N coil undergoes a strain of about 8.7%versus a strain of 18.6% for the Ti-15Mo coil. This means that for agiven amount of exerted energy, the Ti-15Mo wire has about twice theelongation as the MP35N wire. As a result, the Ti-15Mo coil allows formore “stretch”, which is particularly important in applications whereina medical device (e.g., a lead) carrying such a coil may be placed in anarea of a body experiencing movement. A lead carrying a coil formed ofone or more beta titanium alloy filars will stretch with the patient'smovement to enhance patient comfort.

While in the foregoing discussion, the various characteristics of MP35Nare compared specifically to the beta titanium alloy Ti-15Mo, this wasfor illustrative purposes only. A similar comparison may be made betweenMP35N and other beta titanium alloys with similar results. The betatitanium alloys provide important benefits over the MP35N for use inmanufacturing coils for medical devices.

Whereas the foregoing describes the benefits of beta titanium alloys inthe formation of coils, similar and other benefits are achieved whenusing the beta titanium alloys in cable arrangements.

As previously discussed, whereas a coil may be formed using a gatheringor winding process that winds successive turns of the coil around acentral axis, a cable may instead be formed by twisting together wiresthat were previously in a parallel configuration with respect to oneanother. Once so twisted, the cable may be heated to a stress-relievetemperature which will allow the twisted configuration to be retainedonce the “twisting” force has been removed.

Coils and cables exhibit very different properties. For instance, a coilwill generally have a greater elasticity than a cable. As a result, whenforce is exerted on the coil, the coil will stretch to a relatively highpercentage of its total length before permanently deforming or breakingA cable structure will not have the same degree of elasticity. Thus, acoil is able to withstand more longitudinal force than a correspondingcable structure having a same number of wires and being formed of thesame material. On the other hand, the amount of material needed to forma coil is greater than that required to form a corresponding cable, andthe coiling process may be more time-consuming, leading to highermanufacturing costs. Thus, the decision as to whether to utilize a cableor coil structure will be application-specific. In any event, usingbiocompatible beta titanium alloy wires of the type described herein(either with, or without, low-resistance cores) will provide importantbenefits both to coils and cables that are not available with MP35Nwires.

Next, the discussion will turn to various examples of beta titaniumalloy wires arranged in cable configurations.

FIG. 14A is a cross-sectional view of one example of a cable 118according to the current disclosure. The cable 118 includes multiplewires. Each such wire may, in one example, be formed of a tube, or outerlayer, of beta titanium alloy that directly surrounds a low-resistancecore. In this embodiment, seven wires are shown, each including one ofthese tubes 120 a-120 g that may immediately surround a respective oneof the cores 122 a-122 g. The seven wires may be twisted together in amanner to be described below in reference to FIG. 14B.

In the example of FIG. 14A, the wires of cable 118 are surrounded by asingle insulating sheath 126, which may be a polymer that includes, butis not limited to, ethylene tetrafluoroethylene (ETFE),polytetrafluoroethylene (PTFE), silicone rubber or polyurethane. Othermaterials that act as electrical insulators may be used in thealternative. This sheath may be formed in an extruding process thatwraps, or encases, the material around the cable. In another embodiment,a bare cable without sheath 126 may be located within a medical device(e.g., a lead) such that the body of the medical device may itself serveas the insulating sheath.

In another example, rather than be encased by sheath 126 that provides asubstantially uniform layer around the wires of cable 118, the cable mayinstead be coated in a manner that “fills in the gaps” between adjacentwires. For instance, the cable may be fed through a slot or opening of amicro-extruder that applies a thin layer of a polymer such as ETFE tothe entire surface of the cable so that the final cable has apredetermined diameter. The predetermined diameter may be selected sothat the layer has at least some predetermined minimum thickness at itsthinnest point. In one example, this predetermined thickness at itsthinnest point is 0.001 inches. This is discussed further in regards toFIGS. 15A-15D.

In one example, all seven wires within cable 118 may be electricallycoupled one to another. Each of these wires may then be electricallycoupled to the same set of elements to provide redundancy. For instance,each of the wires of cable 118 may be electrically coupled to a sameconducting electrode 38 and a same connector electrode 36 (FIG. 3A). Insuch a scenario, failures may occur in up to six of the wires of cable118 without experiencing an open circuit, so long as one of the wires isstill electrically coupling the two interconnected elements.

A particular embodiment of the foregoing may include seven un-insulatedwires, each having a diameter of between 0.00133 inches-0.00167 inches.The bare 1×7 cable (excluding insulating sheath 126) of such an examplemay have a diameter of between 0.004 inches-0.005 inches.

In another embodiment, one or more of the wires of cable 118 may beprovided with a respective insulating sheath so that these one or morewires are not electrically coupled to at least some of the other wiresincluded in the cable. This may allow some wires to be electricallycoupled to a respective set of elements (e.g., a conducting/connectorelectrode pair) while other ones of the wires within the same bundle areelectrically coupled to a different respective set of elements. In aspecific example, all of the wires may be provided with respectiveinsulating sheaths so that each of the seven wires in cable 118 iscapable of transmitting a different respective signal. Thus, the wiresof cable 118 may be configured in many different ways.

It will be appreciated that any desired degree of redundancy may beprovided by a cable of the type shown in FIG. 14A, with the number ofwires that may be included within the cable being limited by the outerdiameter of the wire and the desired outer diameter of the cable. Insome examples, at least one of the wires, such as the central wire, mayhave a different diameter than the other wires. For instance, a centralwire may have a larger diameter than other surrounding wires. Thus, manycable arrangements are possible within the scope of this disclosure.

FIG. 14B is a side view of a cable according to one embodiment of thedisclosure. This view shows a cable 128, which may be of a type such asshown in FIG. 14A. This cable comprises wires 130 a-130 g, each of whichmay have a low-resistance core and an outer surrounding layer formed ofbeta titanium alloy. As previously discussed, to construct this cable,the wires may be arranged in a substantially parallel configuration withthe ends being held under a predetermined amount of force. A twistingaction is exerted on one or both ends to twist the wires together.

In one example, during manufacturing, each of the multiple wires isretained on a respective spool. An end of each such wire is unwound fromthe spool and threaded into a respective retaining member that retainsthat wire in a predetermined relationship with respect to the otherwires. For instance, to form a cable of the type shown in FIG. 14A, theretention members would retain an end of one wire in a central locationwith six additional retention members retaining six additional wiresaround the central wire. While the retention members are maintained inthis fixed relationship to one another, at least some of the retentionmembers may be rotated as the wires are unwound from the spool to formthe cable. Of course, other mechanisms are available to twist the wires,and this is merely one example.

After being twisted together in this manner, and while the ends of thewires are retained under stress (e.g., to retain the wires in thetwisted configuration), the entire length of the cable may be heated toa stress-relieve temperature that allows the wires to remain in thistwisted configuration after the twisting force is removed. This heatingmay be done by passing the length of the cable through a heated chamberwhile the wires of the cable are still under stress so that each pointin the cable is heated to the stress relieve-temperature. Heating thecable to the stress-relieve temperature changes the physical propertiesof the wires and the cable as a whole, allowing those wires to remaintwisted in a rope-like configuration even after the twisting force hasbeen removed.

As previous discussed, each of the wires in FIGS. 14A and 14B may be ofa type described herein having a beta titanium alloy outer layersurrounding a low-resistance core. One benefit of using this type ofwire in the cabling process as opposed to MP35N involves the superiorcharacteristics of the wire following heating to the stress-relievetemperature. The stress-relieve temperature that is required to form acable of multiple MP35N wires ranges from 500° C.-850° C. Heating theMP35N to such temperatures causes the MP35N wires to become brittle andinflexible. As a result, when subjected to repeated stress (e.g.,bending), the wires may weaken and require replacement. This shortensthe life of the medical device (e.g., lead) that carries the cableformed of the MP35N wire(s).

In contrast to MP35N wires, beta titanium alloy wires do not becomebrittle at their stress-relieve temperature. As one example,stress-relieve heating can be performed at temperatures of between 500°C.-850° C. for under 20 seconds for any of the biocompatible betatitanium alloys. In a particular example, a cable formed of Ti-15Mowires may be heated to a stress-relieve temperature of between 500°C.-650° C. for under 20 seconds without becoming brittle. One even morespecific scenario uses a temperature of between 600° C.-650° C. which ismaintained for under 10 seconds for a Ti-15Mo cable. Heating to atemperature of 625° C. for 9 seconds may be used for this purpose in yetanother embodiment.

In any of the foregoing cases, the physical characteristics of thewires, as well as the cable comprising the wires, will be changed basedon heating to these temperatures for these time periods. For instance,after heating to the foregoing temperatures, the Ti-15No wires carriedby the cable are ductile, and can undergo a high amount of strain. Asimilar result is obtained for any of the other beta titanium alloysdiscussed herein: the wires and resulting cable structures will notbecome embrittled at their respective stress relieve temperatures, whichwill range from between about 500° C.-850° C.

Other benefits similar to those set forth above with respect to coilsare obtained when using beta titanium wires to form cable structures.Beta titanium alloy cables are more biocompatible and do not presentcorrosion issues resulting from metal ion oxidation. Moreover, a cableformed of beta titanium wires will have a much higher fatigue endurancelimit and better kink resistance than a cable that comprises MP35Nwires. As a result, the cables have a longer life. Additionally, suchcables will have a decreased bend radius because of the lower elasticmodulus of the beta titanium alloy as compared to that of MP35N. Thisresults in medical electrical devices (e.g., leads, etc.) that can bemore easily steered into a desired location within the body and whichare more comfortable for the patient. Use of the beta titanium alloywires further allows the cables to be more weld-compatible with othermaterials commonly used in medical devices, including Pt, Pt alloys, Ta,Ta alloys, Nb and Nb alloys.

As another benefit, in a cable embodiment wherein one or more of thebeta titanium alloy wires includes a respective low-resistance core, itis possible to tune the resistance of each such wire. This may beaccomplished by selecting the fraction of the core cross-sectional areato that of the overall wire cross-sectional area, as is discussed abovewith respect to FIG. 7. The cable resistance may be approximated as thein-parallel resistance of the individual wires in one embodiment. Thatis, the cable resistance may be modeled as multiple parallel resistors,with each resistors having a resistance of a respective one of thewires. As such, after tuning the resistances of the individual wires,the cable resistance can be further tuned by selecting the number ofwires to include in the cable. In this manner, the resistance of thecable can be tuned to be similar to, or the same as, a component towhich the cable will be electrically coupled. For instance, the cablecan be tuned to have the same, or a similar resistance, as an electrodeto which it will be coupled, thereby minimizing signal reflections. Thiscan decrease the amount of power needed to send and receive signals.Such tuning is not feasible with silver-cored MP35N wires for reasonspreviously described.

Still other benefits can be obtained by using beta titanium alloys toprovide cable structures that are better suited for MRIconditionally-safe use than silver-cored MP35N. For instance, a higherresistance wire can be provided by selecting such alloys as Ti-15Mo,TLC, and TNCZ instead of MP35N, since these alloys have a higherresistivity than MP35N. In an application wherein the wires comprise aninner core such as shown in FIGS. 14A and 14B, niobium or tantalum maybe selected for use as the core material, since either material exhibitsa significantly higher resistivity than silver. The resulting cable willhave a higher resistance than a cable formed of silver-cored MP35N wire,thereby reducing heating for MRI applications.

Yet another benefit of using beta titanium alloy cables results from thefact that the beta titanium alloy wires 130 a-130 g may be fullyannealed without melting the inner low-resistance cores. The wires aretherefore ductile, further improving the fatigue life of the cable.

Returning again to the cable of FIG. 14B, it may be noted that the cablemay have a “lay” that is similar to the pitch described above withrespect to a coil. This lay is the distance between adjacent twists ofthe same wire, as shown by arrow 132. In one example, the lay will bebetween 0.019 inches to 0.06 inches. In a more specific exampleinvolving a 1×7 cable, the lay may be in the range of between 0.038inches to 0.06 inches. For a 1×3 cable, the lay may be in the range ofbetween about 0.019 inches and 0.04 inches. Thus, the lay may be atleast partially dependent on the number of wires included in the cable.

As may be noted, in the examples of FIGS. 14A and 14B, there is a centerwire (e.g., wires 122 g and 130 g, respectively) within the cable thatmay not be twisted with the rest of the wires, but instead runs alongthe center of the coil. This wire runs substantially along alongitudinal axis of the cable. Because this central wire is not in atwisted configuration, it will not be able to sustain a same level ofstrain as the other wires. In other words, this central wire may havemore susceptibility to bending and longitudinal stress than the otherwires in the cable. Therefore, it may be advantageous in some examplesto eliminate this center wire in favor of a cable configuration in whichevery wire has a twisted configuration. This is discussed further below.

FIGS. 14A and 14B further illustrate a difference between a cable and acoil discussed above. A cable is a solid structure having no centrallongitudinal axis that defines an inner space, or “air core”, runningthe length of the cable. Thus, with a cable design, there is no space inwhich to insert a steering device as is the case in some coil designsdiscussed above. This is discussed below in regards to lead designsusing cable structures.

In the above descriptions of both coils and cables, example wires havebeen described primarily as having low-resistance cores surrounded by alayer of a biocompatible beta titanium alloy, which may be described asa “tube” of this material. In other examples, it may be desirable toutilize biocompatible beta titanium alloy wires that are formedcompletely of that alloy and which omit the low-resistance core. Infact, in embodiments wherein low resistance and/or tuning are lessimportant considerations, there may be no need to include alow-resistance core. A wire without a core may be less expensive tomanufacture, resulting in lower device costs.

In accordance with the foregoing, it should be understood that any ofthe coil or cable structures described herein may be formed of wiresmade completely from the biocompatible beta titanium alloys describedherein but that omit the low-resistance cores. That is, the wires aresolid beta titanium alloy structures to their center, with the corebeing “replaced” instead by beta titanium alloy material. These coil andcable structures will exhibit many of the properties discussed above,having characteristics that are superior to coil or cable structuresformed of MP35N wires. In particular, the benefits listed in Table 3(excluding tunability and minimizing resistance) may be achieved by betatitanium alloy wires regardless of whether low-resistance cores areprovided.

Further in regards to coil configurations, the properties of coils setforth in Table 4 above are achieved by beta titanium alloy wiresregardless of whether those wires have low-resistance cores or are madecompletely of beta titanium alloy. Such properties include the abilityto achieve increased pitch (and hence increased filar count) for a givenouter coil diameter as well as the ability to achieve a smaller coilouter diameter for a coil having the same number of filars. As anotherexample, the benefits achieved by the coil configurations of FIG.10A-10C could largely be obtained from coils formed of wires with, orwithout, the low-resistance, cores (exclusive of tunability andminimizing resistance).

In regards to cable structures, the embrittlement issues related toheating of an MP35N cable to a stress-relieve temperature during acabling process are eliminated by using beta titanium alloy wiresregardless of whether those wires include low-resistance cores.

Because the low-resistance cores are not needed to achieve many of thebenefits described herein, it should be understood that any of the coilor cable embodiments described and/or shown herein may comprise wiresformed completely of biocompatible beta titanium alloy wires thatexclude the low-resistance cores. As such, several specific cableexamples that omit low-resistance cores are next considered forillustration purposes.

FIGS. 15A-15D are cross-sectional views of various cable examples formedof biocompatible beta titanium alloy wires in the manner discussedabove. These wires may be cabled together in a manner similar to thatdescribed above with respect to FIGS. 14A and 14B. In FIG. 15A, threewires are contained within the cable. For the cable of FIG. 15B, four orfive wire may be included in the cable, with an optional fifth wire 141(shown dashed) running down the center of cable. This optional fifthwire 141 has a diameter that is smaller than that of the other fourwires. The cable of FIG. 15C includes four wires in an oblongconfiguration, with one pair of oppositely-situated wires 143 a, 143 bbeing in contact with each other, whereas the other pair ofoppositely-situated wires 145 a, 145 b are not in contact with eachother. This may provide a more stable configuration than a four-wireconfiguration of the type shown in FIG. 15B (that is, when optionalcentral wire 141 is omitted).

Except for the configuration of FIG. 15B that includes optional centerwire 141, the configurations represented by FIGS. 15A-15C do not includea central wire. In such configurations, when cabling is performed in amanner similar to that shown in FIG. 14B, all of the wires in the cablewill be twisted at a similar lay. This may result in a cable havingbetter strain tolerance.

The bare cables (i.e., the cables excluding the insulating layers) ofFIGS. 15A-15C may have diameters of between 003 inches-0.010 inches inone embodiment. In a more specific example, the cables may have an outerdiameter of between 0.004 inches-0.005 inches. When insulating sheathsare applied, the cables of the latter example may have an outer diameterof about 0.005 inches-0.006 inches.

In one example, the three to four larger wires forming the cables ofFIGS. 15A-15C may have diameters of between 0.0010 inches-0025 inches.In a more specific example, the diameters of these wires may be between0.0018 inches-0.0024 inches with the smaller central wire 141 of FIG.15B having a diameter of between about 0.006 inches-0.008 inches. Manyother examples are possible.

FIG. 15D is a cross-section view of yet another type of cable accordingto techniques described herein that includes a central wire 140. Sixintermediate wires are directly adjacent to this central wire 140.Twelve additional wires surround the intermediate wires. As wasdescribed above with respect to FIGS. 14A and 14B, the central wire 140will not be twisted along with the other wires but instead will runsubstantially down the middle of the cable. Since this central wire isnot twisted (and thus has no enhanced capability to “give” or stretchwhen stress is exerted on the cable), the central wire may experience ahigher level of strain than the other wires. For this reason, it may bedesirable to utilize a configuration wherein all wires are twisted, asshown in some examples of FIGS. 15A-15C.

FIG. 15E is a side perspective view of a cable such as shown incross-sectional view in FIG. 15A. This view again shows that there is nocentral wire running the length of the cable, and thus all wires in thecable may be capable of sustaining a same amount of strain.

As is the case described above with respect to FIG. 14B, the cables ofFIGS. 15A-15E have a lay. For a cable such as shown in FIGS. 15A and 15Ehaving three wires, the lay may be between about 0.019 inches and 0.04inches in one example. Other embodiments are possible, however.

In all cases, the cables of FIGS. 15A-15E may be formed according tomethods discussed above with respect to FIGS. 14A and 14B. Inparticular, once the wires are twisted together, the twisted cable maybe heated to a stress-relieve temperature. As previous discussed, onebenefit of using the beta titanium wire in the cabling process asopposed to MP35N wire involves the superior characteristics of the wirefollowing heating to the stress-relieve temperature for cabling. Incontrast to MP35N wires, beta titanium alloy wires, including Ti-15Mowires, do not become brittle at the stress-relieve temperature. As oneexample, the stress-relieve heating for cables formed of beta titaniumalloys, including examples shown in FIG. 15A-15E, can be performed at atemperature between about 500° C.-850° C. In a more specific example,such heating may occur between about 500° C.-650° C. for less than 20seconds for Ti-15Mo wires. An even more specific scenario uses atemperature of 625° C. which is maintained for less than 10 seconds (orapproximately 9 seconds in one particular example) to ensure that aTi-15Mo cable is set in a twisted configuration. Thereafter, the betatitanium alloy wires carried by the cable (and indeed, the cable itself)are ductile, and can undergo a high amount of strain. Thus, in all suchexamples, heating to these temperature ranges for these times changesthe physical attribute of the cables, allowing the cables to retain itsshape even after the twisting force is removed.

Returning to the cross-section views of FIG. 15A-15D, the illustratedcables are shown to have a solid coating surrounding the wires thatdefines a generally circular or oblong profile for these examples. Forinstance, unlike coil 118 of FIG. 14A which carries a sleeve 126 havinga substantially uniform thickness that surrounds the seven wires whileallowing some “open space” to remain between adjacent wires, the cableof FIG. 15A has a solid coating 142. Such a coating may be formed usinga micro-extrusion process that encases the cable within the insulatingmaterial 142. To accomplish this, the length of the cable may be fedthrough a micro-extruder that applies a thin layer of a biocompatiblepolymer such as ETFE to the entire surface of the cable so that thefinal cable has a predetermined shape and size. The layer may be appliedso that at a thinnest point, at least some predetermined minimumthickness is provided, as indicated by arrows 144 of FIG. 15A. In oneexample, this predetermined thickness is 0.001 inches. In otherexamples, any of the embodiments of FIGS. 15A-15E could include aninsulating sheath or sleeve having a uniform thickness, as illustratedin regards to FIG. 14A rather than carrying a solid layer of insulatingmaterial.

In the cable examples of FIGS. 15A-15C having between three and fivewires, the wires may be of a larger diameter than cables of a similardiameter that include more wires, such as shown in the 1×19 cableexample of FIG. 15D. This may be desirable to decrease resistance of thewire in low-frequency applications. However, in higher-frequencyapplications, it may be desirable to include more wires per cable (evenif such wires have a smaller diameter). This is because when higherfrequencies are used, a “skin effect” may be exhibited, with currenttraveling primarily along the surface of the wire. In such cases ahigher-strand cable having smaller wires will exhibit a lower resistancethan a cable with fewer larger wires.

As previously discussed, in cable configurations, all of the wires inthe cable may be electrically coupled one to another. In otherembodiments, one or more of the wires may be provided with a respectiveinsulating layer to insulating the wire electrically from the otherwires in the cable. This may allow the wires that are so insulated tocarry a different signal that than is being carried substantiallysimultaneously by other ones of the wires in the cable.

One or more cables of the type described herein may be carried by animplantable medical device, as discussed in regards to FIGS. 16A-16C.

It may be noted again that all cable configurations of FIGS. 15A-15E,and indeed all cable configurations described herein, could includelow-resistance cores, if desired. Similarly, all coil configurationsdescribed herein may, but need not, include the low-resistance cores.

FIG. 16A is a cross-sectional view of a medical electrical lead of oneembodiment. The lead includes four peripheral lumens 146 a-146 d, eachof which carry a respective cable 147 a-147 d, each of which may be anyof the types described herein or variants thereof. In the particularexample, the cable includes four wires and is configured in a mannersimilar to that shown in FIG. 15B. Each such cable may extend between,and electrically couple, a respective pair of conducting/connectorelectrodes of the type shown in FIG. 3A. In such an example, eachconducting electrode may be used to deliver an electrical signalindependently of the other conducting electrodes. For instance, all fourconducting electrodes 38 (FIG. 3A) could deliver independent signalssubstantially simultaneously to tissue in an example wherein IMD 12 hasfour independent stimulation channels. Alternatively, some or all of theconducting electrodes 38 could be used to sense signals (e.g., whileother electrodes provide stimulation, for instance.)

The embodiment of FIG. 16A further provides a central lumen 148 whichmay receive a guiding device such as a stylet or guide wire used toplace the lead during implantation. In another embodiment, lumen 148could receive a sensing device or another device, such as a fiber fordelivering optical stimulation to tissue, a sensor to sense a signalfrom the tissue, or a device to deliver a pharmacological agent. Inanother embodiment, lumen 148 may be omitted.

Each of the cables shown in FIG. 16A are illustrated as including aninsulating layer that surrounds the four wires of the cable. In such anembodiment, the insulating layer is stripped away from the cable at thelocation wherein the cable is to be electrically and mechanicallycoupled to a respective electrode, such as a conducting electrode 38 ora connector electrode 36 (FIG. 3A). The stripped portion of the cablemay then be mechanically and electrically coupled to the respectiveelectrode by, for instance, a crimping mechanism and/or a spot weld,laser weld, or soldering process.

In another embodiment, one or more of the cables 147 a-147 d need notinclude an insulating layer since the body of the lead provides theinsulation. In this case, stripping of the layer to perform theelectrical coupling to the connector or conducting electrodes need notbe performed.

In yet other embodiments, one or more of cables 147 a-147 d may coupletogether other structures besides connector/conducting electrode pairs.For instance, the lead may carry one or more sensors at the lead distalend which may be coupled via one or more of cables 147 a-147 d toanother structure (e.g., a connector electrode) at the lead proximalend. Any other structure that needs to be electrically coupled to acorresponding structure at the lead proximal end may be so coupledthrough one or more of the cables.

FIG. 16B is another cross-sectional view of a medical electrical leadthat may carry multiple cables. In this case, eight lumens 150 a-150 hare shown, each for receiving a respective cable which may be any of thecables described herein or variants thereof. Thus, in this instance, alead carrying eight conducting electrode pairs could each beelectrically coupled to a respective one of the cables so that eightsignals could be transferred (e.g., sensed or delivered to tissue)substantially simultaneously by the electrodes. As was the case with theexample of FIG. 16A, a lumen 152 is provided to receive a device, whichmay provide navigating/steering capabilities or may be another type ofdevice for sensing, delivering stimulation (electrical or optical),delivering a pharmacological agent, or performing some other task. Inanother embodiment, lumen 152 may be omitted. In such an example whereinthe lumen 152 is omitted, the lead could be placed at a location in abody using a steerable sheath rather than a guiding device positionedwithin the lumen.

FIG. 16C is yet another example of a medical electrical lead thatincludes only a single cable 154 within a central lumen 155. In thiscase, the cable is shown to exclude an insulating layer, with the leadbody 157 itself providing the insulation. This type of configuration maybe desirable if a lead having a very small diameter is required.

This example further shows that the central wire 156 of the cable may belarger than the other wires 154, if desired. This cable may be used toelectrically and mechanically couple a single conductor/connectorelectrode pair. In another example, it may be possible to provideinsulating layers for one or more of the individual wires within thecable so that the wires so insulated could electrically couple adifferent pair of structures (e.g., connector/conducting electrodepairs) than others of the wires. Thus, many configurations are possible.

FIG. 17 is a flow diagram of one method of forming a wire according toone example of the disclosure. First, a tube is formed of abiocompatible beta titanium alloy (180). The tube may have an innerdiameter of between 0.1 inches to 2 inches (or between about 0.25 cm-5.0cm). The beta titanium alloy from which the tube is formed exhibits thebody center cubic (BCC) structure of titanium and may have a modulus ofbetween 30 GPa and 90 GPa.

Next, a core may be formed of material possessing a resistivity of lessthan 25 micro-ohm-cm (182). Silver, which has a resistivity of 1micro-ohm-cm may be used for this purpose. In a different example, thematerial used for this purpose possesses a resistivity of between 10micro-ohm-cm and 20 micro-ohm-cm, and may be a material such astantalum, niobium, or any other biocompatible materials discussed hereinor known in the art that have such a resistivity. This core may beformed by cold-working the core material, or by heating and drawing it.The diameter of the final core is use-dependent and is sized to readilybe inserted within the inner lumen of the beta titanium alloy tube.

In one example, the diameter of the core may optionally be sized so thatafter a wire has been drawn from the tube and core, the wire will have adesired predetermined resistance (183). In particular, the core diametermay be selected so that in the finished wire, the core cross-sectionalarea is a predetermined percentage of the cross-sectional area of thewire as described above with respect to FIG. 7. In this manner, theresistance of the end-product wire is selectable, and may be tuned for aparticular application. For instance, the core may be formed to have arelatively large cross-sectional area if a relativity low resistancewire is desired. Alternatively, a relatively small cross-sectional areamay be selected for the core if the resulting wire is to have a higherresistance. The selected resistance of the wire may be tuned to that ofan interconnecting structure such as a conducting or connectorelectrode.

The core may be inserted into the tube (184), and the core and tube maybe cold drawn (as in drawing it through a die of a predetermined size)to form a wire. (186). This wire may be annealed at the beta transittemperature of the beta titanium alloy and below the melting point ofthe core to obtain a ductile wire (188). This annealing step will changethe physical properties of the wire, allowing the wire to retainductility so that it may optionally be submitted to another cold drawingstep. Because the beta titanium alloy will not become brittle at theannealing temperature, such a wire will not have the embrittlementissues associated with silver-core MP35N wires.

If a wire having a desired outer diameter has been obtained (190),processing may continue to step 192 where a layer of insulating materialmay optionally be applied to the wire. In one example, this involvesdipping the wire in a liquefied ETFE to coat the wire, and then allowingthe insulation material to solidify. Any other biocompatible insulatingmaterial may be used instead as discussed herein, and other processessuch as extrusion may be used to apply this material.

If the desired outer diameter has not been obtained in step 190,processing may return to step 186 wherein the wire is re-drawn throughanother die having yet a smaller diameter and the wire is re-heated asshown in step 188. Steps 186 and 188 may be repeated any number of timesto obtain a wire having a desired diameter.

In one example, the final wire may have a low-resistance core that is,in one embodiment, surrounded by an unbroken layer of the beta titaniumalloy, with the wire having a diameter of between 0.001 inches-0.01inches. Other diameters may be used in other examples.

FIG. 18 is an example flow diagram of a manufacturing process accordingto one specific cable embodiment of the current disclosure. Multiplewires formed of a biocompatible beta titanium alloy wires having amodulus of between 30 GPa and 90 GPa may be obtained, each optionallyhaving a low-resistance core (200). While generally wires used to form asame coil will either all have cores or none will have cores, this neednot be the case. In another example, only some wires need have cores.This may be desirable, for instance, if a lead carries different typesof electrodes, and various filars of a coil are being tuned forconnection to different ones of the electrodes. Optionally, one or moreof the wires may be provided with a respective insulating layer.

Next, the wires may be aligned in a substantially parallel manner withslack removed and ends of the wires being held securely in place (202).A force may be exerted at first ends of the wires, or opposing forcesmay be exerted on both ends of the wires to twist the wires together toform a cable (204). As discussed above, this may be accomplished bythreading ends of wires into retaining members and twisting one or moreof the retaining members as wires are uncoiled from spools, therebyforming the twisted cable. The cable may then be heated to astress-relieve temperature of the beta titanium alloy wires to form aductile cable (206). A particular embodiment heats the cable to betweenabout 500° C.-650° C. for less than 20 seconds for Ti-15Mo wires. Onespecific scenario uses a temperature of 625° C., which is maintained forless than 10 seconds. This heating will change the physical propertiesof the cable, allowing the cable to remain twisted even after thetwisting force is removed.

An outer insulating sheath may be provided (208). For instance, thecable may be dipped in liquefied ETFE. Any other biocompatibleinsulating material may be used for this purpose. Alternatively, anextrusion process may be used to apply the insulating sheath to thecable.

FIG. 19 is a flow diagram of another manufacturing process according toone specific coil embodiment of the current disclosure. Multiple wiresformed of a biocompatible beta titanium alloy wires having a modulus ofbetween 30 GPa and 90 GPa may be obtained, each optionally having alow-resistance core (220). For instance, twelve such wires may beobtained. Optionally, one or more of these wires may be provided with arespective insulating layer (221). The wires may be wound about acentral axis to form a multi-filar coil, with the ratio of the coilpitch to the outer coil diameter being greater than one (222). Inspecific examples, this ratio may be two or three.

In one instance, the wires are wound around a mandrel to form the coil,with the mandrel being removed after winding is completed. In a specificexample, at least some of the multiple filars may be electricallyinsulated one from another or all of the filars may be insulated in thismanner. This may be achieved by providing each of the wires used to formthe coil with a respective insulating coating, such as a coating ofETFE.

Optionally, an insulating sheath may be provided for the coil (224). Inone embodiment, the insulating sheath may be a lead body that carriesthe coil. One or more of the filars of the coil may each be electricallyand mechanically coupled to a different respective element, such as aconducting electrode and/or a connector electrode (226). In a specificembodiment, each of the filars may be coupled to a differentconducting/connector electrode pair to transmit a respective electricalsignal therebetween. Thus, in the specific scenario wherein the coilincludes twelve filars, up to twelve conducting/connector electrodepairs may be so connected to independently transmit twelve signalssimultaneously via the filars of the coil. In some examples, one or morefilars carrying cores may optionally have resistances that are tuned toapproximate or match the resistances of element(s) to which the filar(s)are coupled.

In another example, the inner lumen of the coil defines a space that mayreceive a guiding device such as a stylet, guide wire, or some otherguiding mechanism that can be used to position the coil (and the devicethat carries the coil) within a living body.

Various concepts are described herein. Each concept may be used alone orin conjunction with some, or all, of the other concepts describedherein. As examples, any of the biocompatible beta titanium alloysdescribed herein may be employed to form wires used to construct any ofthe coil configurations or variants thereof described herein. Moreover,one or more such wires may, but need not, include cores. Wires that doinclude cores may, but need not, have resistances that are tuned for agiven application. Similarly, any of the biocompatible beta titaniumalloys described herein may be employed to form wires used to constructany of the cable configurations or variants thereof described herein.Moreover, one or more such wires may, but need not, include cores, andthe wires may optionally have resistances that are tuned.

In some instances, one or more coils and/or one or more cables of any ofthe types described herein may be carried by a same medical device. Forinstance, in a device such as represented by FIG. 16B, a coil of thetype described herein may be carried by central lumen 152, with cablesof any of the types described herein being carried by the peripherallumens 150 a-150 h. Thus, many combinations and adaptations arepossible.

It may further be appreciated that in the methods described herein, someof the steps may be re-ordered within the scope of the disclosure.Moreover, some steps may be omitted entirely. For instance, in somecases, an insulating sheath or layer need not be provided for a wire,coil, or cable, since the body of the medical device (e.g., lead) mayserve this purpose. Similarly, those skilled in the art will recognizethat the filars, wires, and cables described herein may be used tocouple other types of elements besides connectors and electrodes, suchas sensors, or any other type of component that is intended to transmit,receive, or conduct an electrical signal. Moreover, the disclosedembodiments need not be limited to use in medical electrical leads, butmay be used in any other type of medical apparatus carrying suchelements, such as catheters. Thus, the embodiments discussed above aremerely exemplary, with the scope to be defined by the Claims thatfollow.

1. An implantable medical device (IMD), comprising: a coil comprisingmultiple filars, each being formed of a biocompatible beta titaniumalloy having an elastic modulus ranging from 30 GigaPascals (GPa) to 90GPa and comprising at least two elements from a group consisting oftitanium, molybdenum, niobium, tantalum, zirconium, chromium, iron andtin, at least one of the multiple filars being electrically insulatedfrom other ones of the filars; and a structural body carrying the coil.2. The IMD of claim 1, wherein a ratio of pitch of the coil to outerdiameter of the coil is greater than one.
 3. The IMD of claim 2, whereinthe coil comprises at least four filars.
 4. The IMD of claim 3, whereinthe coil has a pitch of greater than 0.03 inches.
 5. The IMD of claim 3,wherein the coil has a pitch of greater than 0.06 inches.
 6. The IMD ofclaim 4, wherein the coil comprises at least eight filars.
 7. The IMD ofclaim 4, wherein the coil comprises twelve filars.
 8. The IMD of claim1, wherein a ratio of pitch of the coil to outer diameter of the coil isgreater than two.
 9. The IMD of claim 1, wherein the coil has an outerdiameter of less than 0.03 inches.
 10. The IMD of claim 1, wherein eachof the filars carries a respective layer of insulating material toelectrically insulate the filar from all others filars.
 11. The IMD ofclaim 10, further comprising multiple elements carried by the structuralbody, each of the elements being electrically coupled to a differentrespective one of the filars.
 12. The IMD of claim 10, furthercomprising at least eight electrodes, each of the at least eightelectrodes being electrically coupled to a different respective one ofthe filars.
 13. The IMD of claim 12, wherein the at least eightelectrodes are segmented electrodes
 14. The IMD of claim 11, wherein atleast one of the filars includes a low-resistance core, wherein a ratioof a cross-sectional area of the core to a cross-sectional area of thefilar is selected to tune a resistance of the filar to a resistance of arespective one of the elements.
 15. The IMD of claim 1, wherein each ofthe multiple filars includes a respective low-resistance core.
 16. TheIMD of claim 15, wherein each of the low-resistance cores is formed of amaterial having a resistivity of less than 25 micro-ohm-cm.
 17. The IMDof claim 16, wherein the low-resistance core is formed of silver,tantalum, a tantalum alloy, niobium, a niobium alloy, platinum, aplatinum alloy, palladium, or a palladium alloy.
 18. The IMD of claim 1,wherein the coil defines an inner space, and further comprising asteering device positioned within the inner space.
 19. A method,comprising: winding a coil of multiple filars, each being formed of abiocompatible beta titanium alloy having an elastic modulus ranging from30 GigaPascals (GPa) to 90 GPa and comprising at least two elements froma group consisting of titanium, molybdenum, niobium, tantalum,zirconium, chromium, iron and tin; and providing at least one of thefilars with an insulating layer to electrically insulate the filar fromthe other filars.
 20. The method of claim 19, wherein a ratio of pitchof the coil to outer diameter of the coil is greater than one.
 21. Themethod of claim 20, wherein the coil comprises at least four filars. 22.The method of claim 21, wherein the coil has a pitch of greater than0.03 inches.
 23. The method of claim 21, wherein the coil has a pitch ofgreater than 0.06 inches.
 24. The method of claim 23, wherein the coilcomprises at least eight filars.
 25. The method of claim 23, wherein thecoil comprises twelve filars.
 26. The method of claim 19, wherein aratio of pitch of the coil to outer diameter of the coil is greater thantwo.
 27. The method of claim 19, wherein the coil has an outer diameterof less than 0.03 inches.
 28. The method of claim 19, further comprisingproviding each of the filars with an insulating layer to electricallyinsulate each filar from all other filars.
 29. The method of claim 28,further comprising: providing at least eight electrodes; andelectrically coupling each of the filars to a different respective oneof the electrodes.
 30. The method of claim 28, further comprising:incorporating the coil into a medical device adapted to perform at leastone of delivering therapy to a patient or sensing a signal from thepatient; and electrically coupling one or more of the filars each to adifferent respective one of multiple elements carried by the medicaldevice.
 31. The method of claim 30, further comprising tuning aresistance of the one or more of the filars to each have substantially asame resistance as a resistance of the respective element to which thefilar is electrically coupled.
 32. The method of claim 31, whereintuning a resistance comprises: providing a low-resistance core;providing a layer of the beta titanium alloy surrounding thelow-resistance core to form a filar; and wherein a ratio of an outerdiameter of the core to an outer diameter of the filar is selected toobtain a selected resistance for the filar.
 33. A medical electricallead, comprising: a lead body; and a coil comprising multiple filars,each being formed of a biocompatible beta titanium alloy having anelastic modulus ranging from 30 GigaPascals (GPa) to 90 GPa andcomprising at least two elements from a group consisting of titanium,molybdenum, niobium, tantalum, zirconium, chromium, iron and tin, atleast one of the multiple filars being electrically insulated from otherones of the filars.
 34. The medical electrical lead of claim 33, furthercomprising at least four electrodes, each being electrically coupled toa different respective one of the multiple filars.
 35. The medicalelectrical lead of claim 33, further comprising at least eightelectrodes, each being electrically coupled to a different respectiveone of the multiple filars.
 36. The medical electrical lead of claim 35,wherein a ratio of pitch of the coil to outer diameter of the coil isgreater than two.
 37. The medical electrical lead of claim 35, whereinthe outer diameter of the coil is less than 0.03 inches.
 38. The medicalelectrical lead of claim 33, wherein the coil comprises twelve filars.39. The medical electrical lead of claim 38, wherein the ratio of pitchof the coil to outer diameter of the coil is at least three.
 40. Themedical electrical lead of claim 33, wherein the coil has more than fourfilars and wherein the bend radius at yield for the coil is less than0.15 inches.